Systems and methods for fault detection in emission-guided radiotherapy

ABSTRACT

Disclosed herein are systems and methods for monitoring calibration of positron emission tomography (PET) systems. In some variations, the systems include an imaging assembly having a gantry comprising a plurality of positron emission detectors. A housing may be coupled to the gantry, and the housing may include a bore and a radiation source holder spaced away from a patient scan region within the bore. A processor may be configured to receive positron emission data from the positron emission detectors and to distinguish the positron emission data from the radiation source holder and from the patient scan region. A fault signal may be generated when the positron emission data from the radiation source holder exceeds one or more threshold parameters or criteria.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.16/100,054, filed Aug. 9, 2018, which claims priority to U.S.Provisional Patent Application No. 62/543,140, filed Aug. 9, 2017, eachof which is hereby incorporated by reference in its entirety.

FIELD

Systems and methods herein relate to patient imaging, which may be usedin diagnostic and/or therapeutic applications, including but not limitedto quality control procedures and fault detection for positron emissiontomography (PET) systems.

BACKGROUND

Positron emission tomography (PET) is a non-invasive imaging techniquethat detects positron annihilation events (e.g., coincidence orcoincident photon events) along a line of response (LOR) using opposingPET detectors. Time of flight (TOF) PET measures a time difference ofcoincidence events at the PET detectors to determine a correspondingannihilation location along the LOR. Determination of an annihilationlocation within a predetermined margin of error is dependent on propercalibration of the detector's time resolution. PET systems, includingTOF PET systems, commonly undergo a daily quality assurance (QA)procedure to verify a time calibration of the PET detectors. Fordiagnostic imaging, a loss of calibration between QA checks may generateinaccurate patient data and require the patient to repeat an imagingsession.

In some applications, emission-guided radiation therapy (EGRT) uses anarray of PET detectors to provide real-time location data of positronemissions originating from a patient tumor and a radiation source totherapeutically irradiate the tumor based on the location data. A lossof calibration of the PET detectors during an EGRT treatment session(e.g., between QA checks) and/or any degradation in the spatialresolution, temporal resolution, energy sensitivity and/or precision, aswell as the inability to precisely determine the location of a patienttumor region relative to the therapeutic radiation source may lead tosuboptimal radiation therapy treatment and damage to healthy tissue.Therefore, it may be desirable to provide real-time fault detection in aTOF PET system that may more quickly identify a time calibration errorand/or faulty PET detectors.

BRIEF SUMMARY

Disclosed herein are systems and methods for emission-guided radiationtherapy using one or more positron emission detectors (PET detectors).Generally, a calibration source may be used to monitor the operation ofthe one or more positron emission detectors. The calibration source maybe a radiation source (e.g., radiation source generating positronannihilation events) that is distinct and spatially separated fromradiation sources located within a patient (e.g., radiotracers and/orimplanted fiducials) and may be held between a plurality of positronemission detectors while treating and/or imaging the patient. Thedetectors may concurrently receive positron emission data from thepatient and calibration source. The positron emission data of thecalibration source may be used to verify the functionality and/orprecision of the positron emission detectors. Although the calibrationsource may be disposed near the patient, the calibration source may beof a size and radioactivity sufficient to be located by the detectorswithout significant risk to the patient.

In some variations, an imaging assembly is provided, comprising a gantrycomprising a plurality of positron emission detectors and a housingcomprising a calibration source holder such as a radiation sourceholder. The housing may be coupled to the gantry. The gantry may furthercomprise a bore and a patient scan or treatment region may be locatedwithin the bore and disposed between the positron emission detectors.The calibration source holder may be stationary and be spaced away fromthe patient scan or treatment region within the bore. The stationarycalibration source holder may be located within the housing or on asurface of the housing. The positron emission detectors may comprise afirst array of rotatable positron emission detectors and a second arrayof positron emission detectors. The assembly may further comprise aprocessor configured to receive positron emission data from the firstand second arrays of rotatable positron emission detectors and todistinguish the positron emission data from the stationary calibrationsource holder and from the patient scan region, and to generate a faultsignal when the positron emission data from the stationary calibrationsource holder exceeds a threshold parameter.

In some variations, the assembly may further comprise a patient support.The patient support may comprise a movable support surface and a base.In some of these variations, the calibration source holder may bedisposed along the surface of the housing at a location above thepatient scan region. In other of these variations, the calibrationsource holder may be located below the movable support surface. In somevariations, a calibration source (e.g., radiation source) may be held bythe calibration source holder. In some of these variations, thecalibration source may comprise a radioactivity of about 1 μCi to 300μCi, e.g., about 2 μCi, about 100 μCi, and an energy of about 511 keV.In other of these variations, the calibration source may comprise ashape with a maximum dimension from about 0.25 inch to about 3 inches,e.g., about 1 inch, about 2 inches. In another variation, the thresholdparameter may be a variability threshold parameter. In yet anothervariation, the processor may be further configured to concurrentlyclassify the positron emission data from the calibration source holderand from the patient scan region. In some variations, the processor maybe configured with a spatial filter to distinguish the positron emissiondata from the stationary calibration source holder and from the patientscan region. In some of these variations, the spatial filter may be useradjustable. In other of these variations, the processor may be furtherconfigured to automatically adjust a geometry of the spatial filterusing a patient treatment plan.

Also described here are other imaging assemblies. In some variations, animaging assembly is provided, comprising a gantry comprising a pluralityof positron emission detectors and a housing comprising a calibrationsource such as a radiation source. The housing may be disposed over thegantry. The gantry may further comprise a bore for a patient to bedisposed between the positron emission detectors. The calibration sourcemay be stationary and spaced away from a patient scan region within thebore. The calibration source may be located within the housing or on asurface of the housing. The positron emission detectors may comprise afirst array of rotatable positron emission detectors and a second arrayof positron emission detectors opposing the first array of detectors.The assembly may further comprise a processor configured to receivepositron emission path data from the first and second arrays ofrotatable positron emission detectors and to classify positron emissionpath data that originates from the stationary calibration source, and togenerate a fault signal when the stationary calibration source positronemission path data exceeds a threshold parameter.

In some variations, a pair of photons emitted by a positron annihilationevent generates a positron emission path. The processor may beconfigured to classify the positron emission path data that originatesfrom the stationary calibration source using a difference between areception time of the pairs of photons within a time threshold parameterrange. In some of these variations, the threshold parameter is alocation deviation threshold. The processor may be configured tocalculate the location of the stationary calibration source based on thereception time difference of the pairs of photons, and to generate thefault signal when the calculated location of the stationary calibrationsource exceeds the location deviation threshold. In other variations,the threshold parameter is a time difference range. The processor may beconfigured to generate the fault signal when a difference between areception time of the pairs of photons is outside of the time differencerange.

In some variations, an imaging assembly is provided, comprising a gantrycomprising a plurality of positron emission detectors and a housingcomprising an annular calibration source such as an annular radiationsource. The housing may be coupled to the gantry. The housing mayfurther comprise a bore and the annular calibration source may be aboutthe bore. The positron emission detectors may comprise a first array ofpositron emission detectors and a second array of positron emissiondetectors opposing the first array of detectors. The assembly mayfurther comprise a processor configured to receive positron emissiondata from the first and second arrays of positron emission detectors andto distinguish the positron emission data from the annular calibrationsource, and to generate a fault signal when the positron emission datafrom the annular calibration source exceeds a threshold parameter.

In some variations, the processor may be further configured toconcurrently classify the positron emission data from the annularcalibration source and from a patient scan region within the bore. Insome of these variations, the processor may be further configured with aspatial filter to distinguish the positron emission data from theannular calibration source and from the patient scan region. In somevariations, the first array and second array of detectors arestationary. In other variations, the first array and second array ofdetectors are rotatable.

In some variations, an imaging assembly is provided, comprising a gantrycomprising a plurality of positron emission detectors. One or morecalibration source holders may be coupled to the gantry such that theone or more calibration source holders are fixed relative to thepositron emission detectors and configured to hold a radiation source.The plurality of positron emission detectors may comprise a first arrayof rotatable positron emission detectors and a second array of rotatablepositron emission detectors opposing the first array of detectors. Aprocessor may be configured to receive positron emission data from thefirst and second arrays of rotatable positron emission detectors and todistinguish the positron emission data from the one or more calibrationsource holders, and to generate a fault signal when the positronemission data from the one or more calibration source holders exceeds athreshold parameter.

In some of these variations, the gantry may comprise a bore. The boremay comprise a patient scan region spaced away from the one or morecalibration source holders. The processor may be further configured todistinguish the positron emission data from the patient scan region inthe bore. In some variations, one or more calibration source holders maycomprise at least four calibration source holders. In another variation,one or more radiation sources may be held by the corresponding one ormore calibration source holders. One or more radiation sources maycomprise a radioactivity of about 1 μCi to 300 μCi, e.g., about 2 μCi,about 100 μCi. In some other variations, one or more calibration sourcesmay comprise a shape selected from the group consisting of a cylinder,sphere, and ring.

Also described here are imaging methods. These methods may comprise thesteps of receiving concurrent positron emission data from a patient anda calibration source spaced away from the patient, using a first arrayof positron emission detectors and a second array of positron emissiondetectors opposing the first array of detectors. The positron emissiondata may be distinguished from the patient and from the calibrationsource. Calibration data may be generated using the positron emissiondata from the calibration source. Patient data may be generated usingthe positron emission data from the patient. A fault signal may begenerated when the calibration data exceeds a threshold parameter.

In some variations, the step of distinguishing the positron emissiondata from the patient and from the calibration source may comprisespatially filtering the positron emission data. In some of thesevariations, a spatial filter may be adjusted before applying the spatialfiltering. For example, a spatial filter may be adjusted based onpatient treatment plan parameters. In some of these variations, thespatial filtering of the positron emission data may comprise excludingthe positron emission data located outside a calibration region and apatient region.

In other variations, receiving the positron emission data from thepatient and the calibration source occurs concurrently with generatingthe fault signal. In another variation, the patient may be treated usinga radiation source concurrently while receiving the positron emissiondata from the patient and from the calibration source. In some of thesevariations, treatment of the patient using the radiation source isstopped in response to generating the fault signal.

In other variations, one or more of the positron emission detectors maybe deactivated based on the generation of the fault signal. In anothervariation, up to three of the first array and second array of detectorsmay be deactivated based on the generation of the fault signal. Thefault signal may comprise a fault in up to three of the detectors. Inyet another variation, all of the detectors may be deactivated based onthe generation of the fault signal. The fault signal may comprise afault in four or more of the detectors. In some variations, one or moreof the positron emission detectors may be calibrated using thecalibration data. In other variations, the positron emission data maycorrespond to lines of response non-intersecting with a patient imagingfield of view of the detectors. The patient imaging field of view maycomprise a patient scan region. In another variation, a fault detectionsystem coupled to the detectors may be verified based on the generationof the fault signal.

One variation of a radiotherapy system (e.g., a radiation treatmentassembly) may comprise a rotatable gantry, a first array of positronemission detectors mounted on the gantry and a second array of positronemission detectors mounted on the gantry opposite the first array ofpositron emission detectors, a therapeutic radiation source mounted onthe rotatable gantry between the first and second arrays of positronemission detectors, a housing disposed over the rotatable gantry andcomprising a bore and a stationary radiation source holder spaced awayfrom a patient region within the bore, and a processor configured toreceive positron emission data detected from the first and second arraysof positron emission detectors. The processor may be configured toextract positron emission data representing positron emission activityoriginating from the stationary radiation source holder, and to generatea fault signal when the extracted positron emission data does notsatisfy one or more threshold criteria. The stationary radiation sourceholder may be located within the housing or on a surface of the housing.The system may further comprise a patient support, the patient supportcomprising a movable support surface and a base. The radiation sourceholder may be disposed along the surface of the housing at a locationabove the patient scan region, e.g., the radiation source holder may belocated below the movable support surface. Some systems may furthercomprise a calibration radiation source held by the radiation sourceholder, the calibration source comprising a radioactivity of about 1 μCito 300 μCi. The calibration radiation source may be configured to beretained by the radiation source holder, the calibration radiationsource comprising a shape with a maximum dimension from about 0.25 inchto about 3 inches (e.g., 1 inch). The calibration radiation source maycomprise a disk-shaped enclosure and a positron-emitting element locatedwithin the enclosure. The processor may be further configured toconcurrently extract the positron emission data representing positronemission activity originating from the radiation source holder and toextract positron emission data representing positron emission activityoriginating from the patient scan region. A threshold criterion maycomprise a spatial filter that selects for positron emission activityoriginating from a location of the stationary radiation source holder. Afault signal may be generated when applying the spatial filter to theextracted positron emission data indicates that the positron emissionactivity does not co-localize with the location of the stationaryradiation source holder. the spatial filter may be user adjustable.Alternatively or additionally the processor may be further configured toautomatically adjust a geometry of the spatial filter using a patienttreatment plan.

In some variations, the first and second arrays of positron emissiondetectors may define an imaging plane, a beam of the therapeuticradiation source may define a treatment plane, and the imaging plane andthe treatment plane may be co-planar. The stationary radiation sourceholder may be co-planar with the imaging plane and the treatment plane.In some variations, the stationary radiation source holder may comprisea groove having a shape that corresponds with a shape of the radiationsource. In some variations, a threshold criterion may comprise athreshold number of coincident photon events detected with a first timedifference (e.g., about 2.5 ns), and the processor may be configured togenerate a plot of an actual number of coincident photon events detectedwith the time difference and a fault signal may be generated when theactual number of coincident photon events occurring with the timedifference does not exceed the threshold number. A threshold positronemission detector criterion may comprise a threshold true-to-randomratio value, where the processor may be configured to generate a ratioof the actual number of coincident photon events occurring within afirst coincidence time window (e.g., from −2.5 ns to +2.5 ns) centeredaround about 0 ns to an actual number of coincident photon eventsoccurring within a second coincidence time window that does not overlapwith the first coincidence time window (e.g., from 17.5 ns to 22.5 ns,not centered around 0 ns, having a similar window width as the firstcoincidence time window) and a fault signal may be generated if theratio does not exceed the threshold true-to-random ratio value. In somevariations, the threshold true-to-random ratio value may be 1 or more,e.g., about 1.1 or more, about 1.3 or more, about 1.5 or more, about 1.6or more, about 2 or more, etc.

In some variations, a threshold criterion may comprise a first expectednumber of coincident photon events to be detected with a first detectiontime difference of about 2.5 ns at a first gantry location of the firstarray of positron emission detectors and a second expected number ofcoincident photon events to be detected with a detection time differenceof about 2.5 ns at a second gantry location of the first array of thepositron emission detectors that is 180° from the first gantry location.The processor may be configured to generate a plot of actual numbers ofcoincident photon events detected within a coincidence time windowbetween −5 ns to +5 ns over a 360° gantry rotation based on positronemission data detected by the first and second arrays of positronemission detectors, and a fault signal may be generated when an actualnumber of coincident photon events detected with a detection timedifference of about 2.5 ns at the first gantry location of the firstarray of the positron emission detectors does not meet or exceed thefirst expected number, and an actual number of coincident photon eventsdetected with a detection time difference of about 2.5 ns at the secondgantry location of the first array of the positron emission detectorsdoes not meet or exceed the second expected number. Alternatively oradditionally, a threshold criterion may comprise an expected number ofcoincident photon events to be detected by each positron emissiondetector of the first and second arrays at each gantry location over a3600 gantry rotation, and the processor may be configured to calculate,using the positron emission data detected by the first and second arrayof positron emission detectors, an actual number of coincident photonevents detected by each positron emission detector of the first andsecond arrays at each gantry location over a 360° gantry rotation, and afault signal may be generated when a difference between the actualnumber of coincident photon events and the expected number of coincidentphoton events exceeds a predetermined difference threshold for at leastone positron emission detector. In some variations, a fault signal maybe generated when the processor does not detect any positron emissiondata representing positron emission activity originating from thestationary radiation source holder. A threshold criterion comprises anenergy resolution spectrum with a coincident 511 keV photon event countabove a peak threshold, and a fault signal may be generated when anenergy resolution spectrum generated from the positron emission datadoes not have a 511 keV photon event count above the peak threshold. Anyof the systems described herein may comprise a display and the processormay be configured to generate a visual indicator and transmitting thevisual indicator to the display. The visual indicator have a firstappearance in the absence of a fault signal and a second appearancedifferent from the first appearance when a fault signal is generated.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A is a block diagram of a variation of a radiation therapyassembly. FIG. 1B is a schematic cross-sectional view of the radiationtherapy assembly depicted in FIG. 1A.

FIGS. 2A-2B are illustrative cross-sectional views of a variation of aradiation therapy assembly.

FIG. 3 is an illustrative cross-sectional view of another variation of aradiation therapy assembly.

FIG. 4 is an illustrative cross-sectional view of yet another variationof a radiation therapy assembly.

FIGS. 5A-5B are illustrative flowcharts of a variation of a method forfault detection.

FIGS. 6A-6B are front and partial cross-sectional side views,respectively, of one variation of a calibration source.

FIG. 7A is an illustrative schematic of one variation of a systemcomprising a calibration source.

FIGS. 7B-7C are elevated perspective and top views, respectively, of onevariation of mount that retains a calibration source.

FIGS. 8A-8C are examples of histogram plots of coincident photon eventcounts over various coincidence time windows.

DETAILED DESCRIPTION

Described herein are radiation therapy and/or imaging systems andmethods for monitoring PET detector parameters and quality metricsduring a radiation therapy treatment session. These systems and methodsmay also be used for calibration separately from a treatment session.Conventional PET detector calibration monitoring is limited to timeperiods between patient radiation therapy treatment and/or imagingsessions. For example, a radiation point source at a known location onthe system may be imaged by PET detectors on an empty couch for a QAprocedure. After a QA procedure has been completed, a patient may beloaded onto the couch and undergo radiation therapy treatment and/orimaging. Conventional QA performed separately from a patient procedurereduces patient throughput and does not monitor the precision and/oraccuracy of PET detectors during a radiation therapy treatment and/orimaging procedure.

Generally, the systems and methods described herein may assist inreal-time monitoring of time calibration for an array of positronemission detectors during a patient image scan and/or radiation therapytreatment. A PET detector that was improperly calibrated or which thecalibration has changed during use and/or otherwise faulty (e.g., due toafterglow effects, malfunctions, etc.) may generate incorrect positronemission location data, which may in turn affect the quality ofradiation therapy treatment and/or patient imaging. A change in amachine parameter or a calibration error may quickly be identifiedsimultaneously during a radiation therapy treatment and/or imagingsession. Because patient workflow does not need to be suspended duringany quality assurance or calibration, this may increase patientthroughput, and may reduce incorrect radiation dose to a patient. Aradiation therapy system as described herein may include an array ofpositron emission detectors (e.g., PET detectors) and a calibrationsource holder for holding a calibration source (e.g., radiation source)at a predetermined (e.g., reference) location. The predetermined,expected location may be compared to an actual computed location of thecalibration source calculated by the system using the positron emissiondetectors. If the locations do not fall within a specified range orwithin a threshold parameter, then one or more of the detectors may beout of calibration and/or faulty, and the system may respond by, forexample, deactivating the detectors and/or halting an imaging and/orradiation therapy treatment based on positron emission data. The systemsand methods may thus provide a safety mechanism to prevent incorrectradiation dose to a patient. In some variations, the imaging systems andmethods described may further comprise a PET detector calibrationmonitoring system configured to monitor other parameters such as atemperature of the PET detector. Generation of a PET detectorcalibration fault signal may be corroborated with the temperature of thePET detector detected by the PET detector calibration monitoring system.In some variations, a discrepancy between the fault signal and thedetected PET detector temperature may indicate a fault in the PETdetector calibration monitoring system.

In some variations, the imaging systems and methods may be used withradiation therapy systems useful for high-energy photon delivery. Aradiation treatment assembly may be useful for emission-guided radiationtherapy, where gamma rays from markers or tracers that are localized topatient tumor regions may be detected and used to direct radiation tothe tumor. Generally, the radiation therapy systems described herein maycomprise a movable gantry, such as a rotating gantry, with positronemission detectors and a radiation treatment source (e.g., MV X-raysource) mounted on the gantry. The positron emission detectors may bemounted on the rotating gantry, and may acquire positron emission data(e.g., emissions from a PET tracer that preferentially accumulates intumor tissue), and the radiation treatment source may deliver aradiation dose to the patient guided by the detector data and atreatment plan. In response to a determination that the positronemission detectors are out of calibration, delivery of further radiationdose may be prevented, thus increasing safety and reducing potentialharm to the patient.

The calibration source (e.g., radiation source) for PET detector timecalibration and/or fault detection may be compact and generateradioactivity sufficient for real-time calibration while minimizingadditional radiation exposure to the patient and/or operator. Forexample, the radiation source may be located between the positronemission detectors (e.g., located in a gantry housing and/or in a boreof the gantry), spatially separated from the patient, and emit enoughpositrons giving rise to coincident photon events to be distinguishableover noise (e.g., cosmic rays). In some variations, the calibrationsource may be located within or on the surface of a (stationary) housingof the rotatable gantry on a top portion and/or a bottom portion of thebore. For example, the calibration source may be located on the surfaceof a housing of the rotatable gantry and co-planar with the positronemission detectors; that is, the positron emission detectors may definean imaging plane along a cross-sectional slice of the bore and thecalibration source may be located on the housing such that it isco-planar with that slice (e.g., at any circumferential location of thebore, such as at the top or 0°, or the bottom or 180°, left side or270°, right side or 90°, etc.). Furthermore, a calibration source holdermay retain the radiation source at a location such that at least some ofthe photons originating from a positron annihilation event may travelalong linear emission paths (e.g., LORs) that do not intersect with thepatient scan or treatment region. Accordingly, calculated locations ofthe calibration source and patient derived from calibration sourceemission data and patient emission data may be spatially separated. Forexample, the location of the calibration source holder (and theradiation source retained within the holder) is not co-localized withthe patient scan or treatment area.

The precision and/or functionality of the positron emission detectorsmay be monitored by comparing a calculated location of the calibrationsource with a reference location or location range of the calibrationsource. For example, calibration data including coincident photonemission time offsets may be compared to reference time offsets tocompute a difference between the calculated location and referencelocation or location range of the calibration source. A differenceexceeding a threshold parameter may generate a fault signal of the PETdetectors. In some variations, fault detection may be performed using astationary calibration source while the positron emission detectors arerotating. In other variations, the calibration source may comprise anannulus shape. In yet other variations, the positron emission detectorsand calibration source may be fixed relative to each other and rotateabout a bore of a gantry. For example, the PET detectors and calibrationsource may be mounted to a rotatable gantry.

I. Systems Radiation Treatment Assembly

Disclosed herein are systems for delivering high-energy photons to aregion of interest (ROI) of a patient while monitoring PET detectorfunction and/or time calibration. FIG. 1A illustrates a block diagram ofa radiation treatment assembly (100) for high-energy photon delivery andreal-time PET detector fault detection. The assembly (100) may include agantry (110) including positron emission (PET) detectors (112), animaging radiation source (114), imaging detector (116), a treatmentradiation source (118), a multi-leaf collimator (120), and a treatmentradiation detector (122). The gantry (110) may be a movable gantry suchas a rotatable gantry that rotates about a longitudinal axis of thegantry (110). For example, a treatment radiation source (118) may bedisposed on a continuously rotatable gantry to generate a radiation beamat one or more gantry angles. In some variations, the gantry (110) maycomprise a ring gantry, and/or may be rotatable about a bore and have anaxis of rotation that is parallel to a longitudinal axis of a bore. Inother variations, the gantry (110) may comprise a C-arm shape. The PETdetectors (112) may comprise any number and configuration to detectpositron emission data (e.g., a pair of 511 keV photons emitted by apositron annihilation event) generated within a patient scan ortreatment region (e.g., within a bore) of the gantry (110). For example,an opposing pair of PET detectors may detect a pair of high-energy 511keV photons and the timing difference between the photon pair may beused to calculate the location of a photon emission origin (i.e.,location of a positron annihilation event) based on time of flight (TOF)of the photons. The positron emission detectors (112) may comprise, forexample, a scintillation detector, comprising one or more of lutetiumorthosilicate (LSO), lutetium-yttrium orthosilicate (LYSO), andlanthanum bromide (LaBr₃). The detectors may be disposed along at leasta portion of a circumference of the gantry (110) and located generallyopposite each other. The positron emission detectors (112) may belocated at the same location along the length of the bore as thetreatment radiation source (118) and multi-leaf collimator (120) (e.g.,along the same tomographic slice). For example, the positron emissiondetectors may define an imaging plane along a cross-sectional slice ofthe bore and the treatment radiation source may be located on the gantrysuch that its irradiation plane or field is co-planar with that slice;that is, the treatment radiation source and the PET detectors may beco-planar (e.g., both mounted on the rotatable ring, arranged such thata beam plane of the treatment radiation source is co-planar with adetection plane of the PET detectors) or may both be located at the samelongitudinal location along the bore (such that the radiation beam planegenerated by the imaging radiation source may be co-planar with theradiation beam plane generated by the therapeutic radiation source).

A rotating mechanism (124) may be coupled to the gantry (110) andconfigured to rotate the gantry (110) from about 10 revolutions perminute (RPM) to about 70 RPM. In some variations, the rotating mechanism(124) may rotate the gantry (110) such that the detectors (112), imagingradiation source (114), imaging detector (116), treatment radiationsource (118), multi-leaf collimator (120), and treatment radiationdetector (122) may rotate about a rotational axis of the gantry (110).In some variations, the detectors (112) may rotate about the gantry(110) while in other variations the detectors (112) may be stationary.

The imaging radiation source (114) and a corresponding imaging detector(116) may be used to generate patient image data (e.g., CT images, MRimages), and in some variations may comprise a kV source and kVdetector. The patient image data may be used to register the patient(e.g., identify the patient's location with respect to the radiationtreatment assembly components) and/or aid delivery of treatmentradiation delivery to the patient. The treatment radiation source (118)may deliver a treatment radiation dose to the patient in a bore of thegantry and may comprise, for example, a linear accelerator (linac) and amagnetron (e.g., MV X-ray source). The treatment radiation beam may beshaped by a beam-shaping assembly coupled to the treatment radiationsource (118) to deliver a prescribed radiation dose to the ROI using aplurality of radiation beams output from a plurality of gantry angles.For example, the beam assembly may comprise a multi-leaf collimator(120) coupled to the treatment radiation source (118) and may be locatedin a treatment radiation beam path for shaping the treatment radiationbeam delivered to the patient. The multi-leaf collimator (120) maycomprise a plurality of leaves and corresponding actuation mechanismsconfigured to independently move (e.g., open and close) the leaves inone or more axes (e.g., X-axis, Y-axis). For example, the multi-leafcollimator (120) may be a binary multi-leaf collimator. The treatmentradiation detector (122) (e.g., MV detector) may oppose the treatmentradiation source (118). The treatment radiation detector may be locatedalong the treatment radiation beam path and may acquire treatmentradiation data. The positron emission detectors (112) may be arrangedsuch that they are not in the treatment radiation beam path. Thetreatment radiation source (118) may generate any type of ionizingradiation, for example, photon radiation (e.g., X-rays and gamma rays)and/or particle radiation (e.g., electrons, protons, neutrons, carbonions, alpha particles, and beta particles). In some variations, theimaging radiation source (114) and treatment radiation source (118) mayhave separate components (e.g., linac, beam converter assembly) while inother variations the sources (114, 118) may share one or more components(e.g., share the same beam converter assembly).

The assembly (100) may further include a housing (130) configured tohold a calibration source (132), a processor (140), memory (142), and apatient support (150). FIG. 1B is a non-limiting schematic example ofthe radiation treatment assembly (100) depicted in FIG. 1A wherecalibration source (132) is not illustrated in FIG. 1B. The housing(130) may enclose the gantry (110) and provide a barrier between thepatient (160) and the gantry (110). For example, the housing (130) maybe coupled to the gantry (110) and provided between the patient support(150) (e.g., couch) and the positron emission detectors (112). Thegantry (110) may rotate while the housing (130) remains stationary. Thehousing (130) may comprise a bore (134) or opening in which the patient(160) and patient support (150) may be disposed. For example, a patient(160) disposed on the patient support (150) may be moved in and out ofthe bore (134) of the housing (130).

The calibration source (132) may be a radiation source configured togenerate radiation sufficient for the detectors (112) to locate thecalibration source in real-time (e.g., during a patient image scanand/or patient treatment session) without exposing the patient and/oroperator to significant additional radiation. In some variations, thecalibration source may comprise a positron emitting material comprisingone or more isotopes such as 22-Na, 68-Ge, 68-Ga, and the like. Emittedpositrons may collide with electrons in an annihilation event togenerate gamma rays (e.g., a pair of diametrically opposed photons) thattravel along a linear path (e.g., line of response or LOR). Detectedphoton pairs are classified as a coincidence event if they are detectedby opposing positron emission detectors (112) within a pre-determinedtime window (e.g., coincidence time window). The detectors record thedetection location and reception time. A reception time differencebetween a pair of coincidence photons is referred to as time of flight(TOF) and may be used to determine the origin of the positronannihilation event along the LOR. The TOF measurement exhibitsuncertainty and corresponds to a timing resolution of the detectors(112). This uncertainty in the positron annihilation event and timingresolution may be represented or characterized by a probabilitydistribution (e.g., Gaussian distribution) or related parameter, whichmay be further characterized by a Full Width at Half Maximum (FWHM) ofthe Gaussian distribution of the location derived from TOF measurement.

In some variations, the calibration source (132) may comprise aradioactivity of about 1 μCi to 300 μCi, e.g., about 2 μCi, about 100μCi. Accordingly, the calibration source (132) may emit enough positronsper second (e.g., annihilation events) for the positron emissiondetectors (112) to receive positron emission data allowing the processor(140) to monitor PET detector calibration (e.g., distinguish positronemissions of the calibration source (132) from the emissions of thepatient (160) using time offset data). The rate of positron-emission ofthe calibration source may be known and the emission rate (andoptionally, the positron annihilation rate) may be used by thecontroller processor to determine whether the positron emissiondetectors are faulty and/or calibrated properly. For example, the LORdetection rate as measured by the positron emission detectors may becompared with the known positron-emission rate of the calibrationsource. If the LOR detection rate (i.e., the LORs and/or coincidentphoton events that may be attributed to the calibration source becausethe LORs intersect the known location of the calibration source and/orcalibration source holder) is greater than or less than the knownpositron-emission rate (and/or an expected LOR emission rate calculatedbased on the known positron-emission rate) by specified tolerancethresholds, the processor may generate a notification to the userindicating that the LOR detection rate differs from the expected rate.Optionally, the controller processor may generate an interlock signalthat pauses or ceases treatment radiation delivery until the user canverify that the positron emission detectors are functioning properlyand/or calibrated.

A shape of the calibration source (132) is not particularly limited andmay comprise any geometric shape such as a cylinder, sphere, ring, rod,disc, line source, etc. In one variation, a calibration source maycomprise a housing or enclosure and a radioactive (e.g.,positron-emitting) element located within the enclosure. The housing orenclosure may be disk-shaped and/or made of a non-radioactive or inertmaterial, such as Mylar, Teflon, epoxy, and/or glass. The radioactiveelement may be embedded within the housing or enclosure. In somevariations, the radioactive element may be a pellet, bead, seed,capsule, droplet, gel, etc., The calibration source (132) may beoriented in any direction so long as the calibration source (132) islocated at the same location along the length of the gantry (110) as thePET detectors (112), i.e., co-planar with the PET detectors. Forexample, a rod shaped or line source shaped calibration source may bearranged in parallel with a longitudinal axis of the gantry (110). Insome variations, the calibration source may (132) comprise a shape witha maximum dimension from about 0.25 inch to about 3 inches, e.g., about1 inch, about 2 inches. In some variations, the calibration source (132)may comprise one or more positron-emitting capsules that each contain aquantity of positron-emitting tracer(s), where each capsule has amaximum dimension of about 2 cm. Some capsules may have a maximumdimension of no more than about 300 μm. For example, a calibrationsource may comprise a disk-shape enclosure with a diameter of about oneinch, a thickness of about 0.25 inch, and a radioactive capsule with adiameter of about 0.039 inch (e.g., about 1 mm). The radioactive capsulemay be embedded within an epoxy well in the enclosure, about halfwaythrough the thickness of the disk and at the center of the disk.Positron-emitting capsules or calibration sources that are relativelysmall (e.g., less than about 2 cm, less than about 1000 μm, less thanabout 500 μm, less than about 300 μm, etc.) may be more easily containedor isolated (to prevent unwanted contamination) and may have arelatively longer half-life (e.g., about 2 years or more, about 2.6years). The calibration source may comprise an array ofpositron-emitting capsules, arranged in a linear configuration and/ordistributed radially about the bore (134). For example, a calibrationsource may be located at the top of a bore (e.g., at 0°), bottom of abore (e.g., at 180°), or any radial or angular position about the bore(e.g., at 90°, 270°, 30°, 120°, 600, 150°, 200°, 300°, etc.). In somevariations, a plurality of calibration source may be located at radiallyand/or bilaterally symmetric locations about the bore (e.g., foursources at 0°, 90°, 180°, and 270°; two sources at 0° and 180°; foursource at 30°, 150°, 210°, 330°, etc.). A calibration source (132) maycomprise a first positron-emitting capsule at a first location about abore or patient area, and a second positron-emitting capsule at a secondlocation across from (e.g., about 180 degrees from) the firstpositron-emitting capsule. Alternatively or additionally, a firstpositron-emitting capsule may be located at a first end of a first arrayof PET detectors, a second positron-emitting capsule located at a secondend of the first array of PET detectors, a third positron-emittingcapsule located at a first end of a second array of PET detectors, and afourth positron-emitting capsule located at a second end of the secondarray of PET detectors. More generally, the positron-emitting capsulesof a calibration source may be located outside of a patient area or thebore of a gantry. The calibration source (132) or positron-emittingcapsule may comprise a radioactive portion and a non-radioactive housingenclosing the radioactive portion. In one variation, the non-radioactivehousing may be disc shaped and a radioactive portion may be sphericaland located at a center of the disc. In another variation, a ring shapedradiation portion may be disposed in a disc shaped non-radioactivehousing. In yet another variation, the non-radioactive housing may becylindrical and a radioactive portion may be disposed in spaced apartwells located, for example, at the ends of the cylinder.

In some variations, the patient support (150) (e.g., couch) may comprisea support surface and a base (not shown) for control of positioning of apatient in the assembly (100). The base may be fixed to the ground andthe support surface may be coupled to the base such that the supportsurface may move in and out of a bore of the gantry (110). The patientmay be disposed on the support surface to be imaged and/or treated bythe assembly (100) (e.g., the patient lying flat on the patient support(150)).

The processor (140) may incorporate data received from the memory (142)and positron emission detectors (112) to compute a location of thecalibration source (132) based on detected emission data. Based on thecalculated location, positron emission data may be classified asoriginating from one of the calibration source and the patient scanregion. The memory (142) may further store instructions to cause theprocessor (140) to execute modules, processes and/or functionsassociated with the system (100), such as fault detection and safety(e.g., deactivation of one or more system components, stopping radiationtherapy treatment, outputting system status, etc.). For example, thememory (142) may be configured to store location data of the calibrationsource (132), one or more threshold parameters, a patient treatmentplan, one or more spatial filters, positron emission data (e.g.,positron emission path data), calibration data, patient data, andoperator input.

Memory (142) may store a location of the radioactive source held in theradioactive source holder. The processor (140) may be any suitableprocessing device configured to run and/or execute a set of instructionsor code. The processor may be, for example, a general purpose processor,a Field Programmable Gate Array (FPGA), an Application SpecificIntegrated Circuit (ASIC), a Digital Signal Processor (DSP), and/or thelike. The processor may be configured to run and/or execute applicationprocesses and/or other modules, processes and/or functions associatedwith the system and/or a network associated therewith (not shown). Theunderlying device technologies may be provided in a variety of componenttypes, e.g., metal-oxide semiconductor field-effect transistor (MOSFET)technologies like complementary metal-oxide semiconductor (CMOS),bipolar technologies like emitter-coupled logic (ECL), polymertechnologies (e.g., silicon-conjugated polymer and metal-conjugatedpolymer-metal structures), mixed analog and digital, and/or the like.

The memory (142) may include a database (not shown) and may be, forexample, a random access memory (RAM), dynamic random access memory(DRAM), static random access memory (SRAM), a memory buffer, a hard diskdrive, optical disc, magnetic tape, an erasable programmable read-onlymemory (EPROM), an electrically erasable read-only memory (EEPROM), aread-only memory (ROM), Flash memory, solid state drive (SSD), memorycard, etc. The memory (142) may store instructions to cause theprocessor (140) to execute modules, processes and/or functionsassociated with the system (100), such as fault detection.

The system (100) may be in communication with other devices (not shown)via, for example, one or more networks, each of which may be any type ofnetwork. A wireless network may refer to any type of digital networkthat is not connected by cables of any kind. Examples of wirelesscommunication in a wireless network include, but are not limited tocellular, radio, and microwave communication. However, a wirelessnetwork may connect to a wireline network in order to interface with theInternet, other carrier voice and data networks, business networks, andpersonal networks. A wireline network is typically carried over coppertwisted pair, coaxial cable or fiber optic cables. There are manydifferent types of wireline networks including, wide area networks(WAN), metropolitan area networks (MAN), local area networks (LAN),Internet area networks (IAN), campus area networks (CAN), global areanetworks (GAN), like the Internet, and virtual private networks (VPN).Hereinafter, network refers to any combination of combined wireless,wireline, public and private data networks that are typicallyinterconnected through the Internet, to provide a unified networking andinformation access solution. The system may be configured to provide notonly patient diagnostic or therapeutic data, but also machinecalibration data (e.g., calibration data) and QC data to the patient'selectronic healthcare record and/or to electronic record systems used byaccreditation agencies, such as the American College of Radiology, theJoint Commission, the UK Accreditation Services (UKAS), and the EuropeanAssociation of Nuclear Medicine's EARL program, for example.

Calibration Source

As described in detail below, a calibration source of a radiationtherapy assembly may comprise a number of configurations and/orlocations relative to the assembly. As used herein, a calibration sourcemay be a radiation source comprising a substance that emits positronswithin a field of view of the PET detectors sufficient to monitor adetector time calibration while minimizing dose to the patient. Forexample, the radioactive source described with respect to FIGS. 2-4 mayemit photons with an energy of 511 keV for each positron annihilationevent. In other variations, the radioactive source of a calibrationsource may comprise a substance that emits radiation having a differentenergy level. The calibration sources may be located on a radiationtreatment assembly such that at least some lines of response originatingfrom the calibration source do not intersect a patient and/or patientsupport disposed in a bore of the gantry. That is to say, thecalibration source may intersect at least one line provided betweenopposing positron emission detectors and which are unobstructed by apatient and/or patient support (although the line may pass through otherstructures such as the gantry housing). As an illustrative example, thecalibration source may be disposed along a housing coupled to the gantryat a location at least about 20 cm above the patient support. Spatialseparation between the calibration source and the patient may reduceerrors in a calculated location of the calibration source and thepatient. A fault signal for one or more positron emission detectors maybe generated using the positron emission data of the calibration source.The system may respond appropriately to the fault signal (e.g., output adetector status, deactivate faulty detectors, and/or stop treatment).

One variation of a calibration radiation source is depicted in FIGS.6A-6B. As depicted there, calibration radiation source (600) maycomprise a disk-shaped enclosure or housing (602) and apositron-emitting element (604) located within the enclosure, at thecenter of the disk. The enclosure (602) may be made of a non-radioactivematerial and may have a well (606) within which the positron-emittingelement (604) may be located. In some variations, the enclosure (602)may be made as a solid disk, a well (606) may be created within the disk(e.g., optionally, located on the center of the disk as represented bythe dotted line (607)), the positron-emitting element (604) may beinserted into the well (606), and the well may then be filled with anon-radioactive material, such as an epoxy. The diameter (601) of theenclosure (602) may be from about 0.25 inch to about 3 inch, e.g., about1 inch, about 2 inches. The thickness (603) may be from about 0.1 inchto about 0.5 inch, e.g., about 0.25 inch. The size of thepositron-emitting element (e.g., capsule) may be from about 150 μm toabout 500 μm, e.g., about 200 μm, about 250 μm, about 300 μm, etc. Thepositron-emitting element (604) may be inserted or embedded such that itis halfway between the thickness of the disk. For example, thepositron-emitting element (604) may be located at a distance (608) fromone side surface of the enclosure (602), where the distance (608) may befrom about 0.01 inch to about 0.4 inch, e.g., about 0.12 inch, about 0.2inch, etc.

It should be understood that the systems described below do not requireeach of the components depicted above in FIG. 1. For example, atreatment radiation source (e.g., MV X-ray source) may not be includedin the system variations depicted in FIGS. 2-4. The systems depicted inFIGS. 2-4 may be imaging systems that do not have a treatment radiationsource, or they may be radiation treatment systems with a treatmentradiation source that has been omitted for ease of explanation.

A. Stationary Calibration Source

Variations of an imaging assembly described here may comprise aplurality of positron emission detectors and a stationary calibrationsource for real-time PET detector fault detection. FIGS. 2A-2B arecross-sectional schematic views of a radiation treatment assembly (200)including a rotatable gantry (210) having mounted thereon a first arrayof positron emission detectors (212) and a second array of positronemission detectors (214) opposing the first array of detectors (212).Each array of detectors (212, 214) may comprise a plurality of positronemission detectors. A housing (220) may couple to the gantry (210)(e.g., the housing (220) may be disposed over the gantry (210)), and thehousing (220) may comprise a bore (224) in which a patient support (250)may be disposed. Each positron emission detector has an imaging field ofview that is the angle through which that detector is sensitive topositron emissions. As used herein, a patient scan region (242) (e.g.,patient imaging field of view) may be represented by a volume of apatient (240) and/or patient support (250). That is, the contours of thepatient (240) and/or the patient support (250) detectable by thepositron emission detectors may define a patient scan region (242). Thepatient support (250) may comprise a movable support surface on whichthe patient (240) may be disposed. The movable support surface may becoupled to a base (not shown).

The housing (220) may comprise a stationary calibration source holder(222) and/or calibration source (230) (e.g., radiation source) spacedaway from the patient scan region (242) within the bore (224) and withina field of view of the positron emission detectors (212, 214). Forexample, the calibration source holder (222) and/or calibration source(230) may be located on a surface of the housing (220) (e.g., facing thepatient support (250)). For the sake of illustration, the calibrationsource holders (222) and/or calibration sources (230) of FIGS. 2A-2Bproject from the housing (220) towards the patient (240) although othervariations are contemplated. For example, the calibration source holder(222) and/or calibration source (230) may be located within a recess orwithin the housing (220) between the first array and second array ofdetectors (212, 214). The assembly (200) may comprise a singlecalibration source holder (222) and calibration source (230) where FIGS.2A-2B illustrate two calibration source holders (222) and calibrationsources (230) for ease of explanation, however, it should be understoodthat in some variations, there may be a single calibration source holderand calibration source at either of the locations depicted in thefigures (e.g., only at the top or only at the bottom portion of thebore), or additional calibration source holders and calibration sourcesat additional locations around the bore. The calibration source holder(222) may be configured to hold any of the calibration sources describedherein (e.g., a positron emitting radiation source). The calibrationsource holder (222) may securely hold the calibration source (230) at adesired location (e.g., reference location or location range) using anyknown means, including but not limited to adhesives, a closuremechanism, and an interference fit or mechanical interfit between thecalibration source holder (222) and calibration source (230). A closuremechanism may include a cap, cover, lid, plug, latch, lock, threadedring, and combinations thereof. The calibration source holder (222) mayallow the calibration source (230) to be removably coupled to thehousing (220) such that the calibration source (230) may be replacedafter its useful lifespan.

In some variations, the calibration source holder (222) may be disposedalong the surface of the housing (220) at a location furthest from thepatient scan region (242). In one example, the calibration source holder(222) may be disposed at a top, center location of the housing (220)(e.g., where the calibration source holder (222) is disposed). Thislocation may be the furthest away from the patient (240) and thereforeprovide the least amount of additional dose from the calibration source(230) to the patient (240). In particular, a patient support (250) maybe disposed below a rotational axis of the gantry (210) (e.g.,mechanical isocenter axis) such that the patient (240) on the patientsupport (250) intersects a rotational axis of the gantry (210).Therefore, a distance from a top, center location of the housing (220)to the patient (240) may be greater than the distance of any other pointalong the surface of the housing (220).

In some variations, the calibration source holder (222) may be disposedalong the surface of the housing (220) above or below the supportsurface of the patient support (250). The calibration source holder(222) may be disposed above a horizontal plane of the patient support(250) and/or perpendicular to the patient support (250). In some ofthese variations, the calibration source holder (222) may be disposedabove the patient scan region (242). For example, the calibration sourceholder may be disposed at least about 20 cm above the support surface ofthe patient support (250). In other words, the calibration source holder(222) disposed along the surface of the housing (220) may be above thehighest point of the patient (240). In other variations, the calibrationsource holder (222) may be disposed below a horizontal plane of thepatient support (250) and/or perpendicular to the patient support (250).For example, as shown in FIGS. 2A-2B, the calibration source holder(222) may be disposed at a bottom, center location of the bore (224).

FIG. 2B illustrates the first array and second array of detectors (212,214) at a second position rotated relative to the detectors (212, 214)at a first position in FIG. 2A. The positron emission detectors (212,214) detect photon pairs emitted by the calibration source (230) andpatient (240). The photon pair travels along a line of response (232)originating from the calibration source (230) and extending towards thefirst array of detectors (212) and the second array of detectors (214).At least some of the calibration source LORs (232) are non-intersectingwith the patient scan region (242) (e.g., patient imaging field of viewof the detectors including patient (240) and/or patient support (250))such that at least some of the calibration source LORs (232) arespatially separated from patient lines of response (234) emitted fromthe patient (240). Consequently, the calibration source holder (222) islocated for a calibration source (230) to form at least one LOR (232)unobstructed by the patient (240) and/or the patient support (250).Although not illustrated in FIGS. 2A-2B, the gantry (210) may furthercomprise a treatment radiation source and treatment detector between thefirst array and second array of detectors (212, 214) having a treatmentfield of view (260).

In some variations, a stationary calibration radiation source holder maycomprise a groove or recess within an internal wall of the bore, or maybe a groove or recess within a structure that may be attached to aninternal wall of the bore. Optionally, a cover may be removably disposedover the groove or recess to retain the calibration source within thegroove or recess, and engaged over the groove or recess via anyattachment mechanism (e.g., snap-fit, friction-fit, screw-fit, with orwithout the use of additional one or more screws or fasteners). When thecalibration source is to be replaced, the optional cover may be removed.

FIG. 7A depicts one variation of a radiation treatment assembly orimaging system (700) comprising a bore (702) and a patient platform(704) movable along a longitudinal axis of the bore (702). The system(700) may comprise a rotatable gantry (not shown) that may have an axisof rotation that is collinear with the central axis of the bore (702),and a first array of positron emission detectors (706 a) mounted on thegantry and a second array of positron emission detectors (706 b) mountedon the gantry across from (i.e., opposite to) the first array ofpositron emission detectors (706 a). The first and second array ofpositron emission detectors may be located along a plane or slice (708)of the bore (e.g., orthogonal to the longitudinal axis of the bore). Invariations where the system (700) is a radiation treatment assemblycomprising a treatment radiation source mounted on the rotatable gantrybetween the first and second positron emission detector arrays, thetreatment beam plane may be co-planar with the positron emissiondetector slice (708). In this variation, a calibration radiation source(710) and the calibration radiation source holder (714) may be locatedwithin the bore (702), at a bottom location (701 b) of the bore, belowthe patient support (704). Alternatively or additionally, calibrationradiation sources and/or holders may be located at one or morecircumferential locations about bore (e.g., at 701 a or 0°, at 701 b or180°, at 701 c or 90°, and/or at 701 d or 270°, above the patientsupport, etc.). The calibration source holder (714) may be a groove orrecess within an internal surface of the bore (702), with or without acover, as described above. In some variations, a calibration sourceholder located above the bottom region of a bore may comprise a cover toprevent the calibration source from falling out, while a calibrationsource holder located at the bottom of the bore may not have a cover.Where a calibration source (710) has a disk-shape, the calibrationsource holder (714) may also have a corresponding disk-shape.

In some variations, the calibration radiation source holder may beintegral with, and/or a part of, a housing of other components of thesystem. For example, any of the systems described herein may compriseone or more optical cameras and/or light sources within a bore, andcomprise a camera and/or lighting mount attached to the internal wall ofthe bore. In some variations, the camera and/or lighting mount maycomprise a groove or recess that is configured to be retain acalibration source (i.e., configured to be a calibration radiationsource). For example, the camera and/or lighting mount may comprise arecess that has a size and shape that corresponds with the size andshape of a calibration source and an optional calibration sourceengagement structure, such as a cover (such as any of the coversdescribed above), clasp, and/or magnetic source holder. Optionally, thecalibration radiation source holder and/or mount may have one or morealignment structures, such as one or more grooves and/or protrusionsthat may help to retain the calibration radiation source in precisealignment with the other components of the radiation treatment assembly.For example, the holder and/or mount may have one or more protrusionsthat restrict motion of the calibration source in a first direction(e.g., IEC-Y) and/or one or more additional protrusions that restrictionmotion of the calibration source in a second direction (e.g., IEC-X). Insome variations, one or more protrusions may include a wall portion of acamera and/or lighting mount. Alternatively or additionally, a groove orslot within which the calibration source may be seated may restrictmotion of the calibration source. In some variations, the holder and/ormount may be centered or aligned with an isocenter of the system (e.g.,along a central longitudinal axis of the bore).

FIGS. 7B-7C depict an elevated perspective view and a top view,respectively, of a lighting mount that may be attached to an internalsurface or wall of the bore, the lighting mount comprising a calibrationsource holder. One or more light sources, such as LEDs, may be attachedto the internal surface of the bore via the lighting mount. A lightingcover having a corresponding shape and size (e.g., footprint) as thelight mount may be disposed over the mount. Turning to FIGS. 7B and 7C,the lighting mount (700) may comprise a walled tray or enclosure (702)comprising one or more bore attachment structures (704), a calibrationsource holder alignment protrusion (706) and a calibration sourcereceiving portion or holder (708). The calibration source holder orreceiving portion (708) for a calibration source shaped as a circulardisk may comprise a circular recess that has a diameter that is slightlylarger than the diameter of the circular disk so that the calibrationsource may be retained within the recess. The bore attachment structure(704) may comprise a bracket that engages a corresponding notch on theinternal wall of the bore, and/or may comprise structures for ascrew-fit or any other such engagement with the bore. Optionally, thelighting mount (720) may also comprise one or more cover attachmentstructures (710), which may comprise a clip or clamp that may be used toretain or grasp a corresponding protrusion, flap, or tongue of anoptional lighting cover (not shown). In the variation depicted in FIGS.7B-7C, the lighting mount (720) may have an elongated oblong shape,where the calibration source holder or receiving portion (706) may belocated at one end of the elongated shape. For example, the walledenclosure (702) may comprise a main oblong portion where the one or morelight sources may be attached and a narrow extension to the main oblongportion where the calibration source holder or receiving portion may belocated. This may help to segregate the calibration radiation sourcefrom the light sources, which may help reduce radiation damage to thelight sources.

While the variations described herein may be directed to a system with asingle calibration radiation source and/or a single calibrationradiation source holder, it should be understood that similar structuresand features may apply to systems with a plurality of calibrationradiation sources and/or calibration radiation source holders. Forexample, a plurality of calibration radiation sources and/or holders maybe distributed circumferentially about the bore, at a plurality of boreangles (e.g., bore angle locations 701 a-701 d, as well as anglesbetween).

B. Annular Calibration Source

FIG. 3 is a cross-sectional schematic view of a radiation treatmentassembly (300) that may be configured to provide real-time calibrationmonitoring of PET detectors and include a gantry (310) having a firstarray of positron emission detectors (312) and a second array ofpositron emission detectors (314) opposing the first array of detectors(312). Each array of detectors (312, 314) may comprise a plurality ofpositron emission detectors and mounted to a movable or stationarygantry. Accordingly, the PET detector arrays may rotate or bestationary. A housing (320) may couple to the gantry (310) (e.g., thehousing (320) may be disposed over the gantry (310)), and the housing(320) may comprise a bore (322) in which a patient (340) and a patientsupport (350) may be disposed. A patient scan region (342) (e.g.,patient imaging field of view) may be represented by a volume of thepatient (340) and/or patient support (350). That is, the contours of thepatient (340) and/or the patient support (350) detectable by thepositron emission detectors may define a patient scan region (342). Thepatient support (350) may comprise a movable support surface on whichthe patient (340) may be disposed and the movable support surface may becoupled to a base (not shown).

The housing (320) may further comprise an annular calibration source(330) (e.g., annular radiation source) about at least a portion of thebore (322). The radioactivity level of an annular calibration source maybe similar to the radioactivity level of any of the calibration sourcesdescribed herein, e.g., about 1 μCi to 300 μCi, e.g., about 2 μCi, about100 μCi. For example, the annular calibration source (330) may bestationary and located below a surface of the housing (320) and/or maycomprise the surface of the housing (320). The annular calibrationsource (330) may be located between the first array and second array ofdetectors (312, 314). As shown in FIG. 3, the annular calibration source(330) may comprise a continuous annulus. Regardless of the configurationof the positron emission detectors (312, 314) (e.g., rotatable orstationary), the annular calibration source (330) may be located withina field of view of the positron emission detectors (312, 314).

The positron emission detectors (312, 314) detect photons emitted by theannular calibration source (330) and patient (340). FIG. 3 illustratesan annular calibration source LOR (332) intersecting the annularcalibration source (330) at a first location (334) and a second location(336). The LOR (332) may be unobstructed by the patient (340) and/or thepatient support (350). The positron annihilation event corresponding tothe LOR (332) may originate from either the first or second location(334, 334) of the annular calibration source (330). A TOF PET system mayuse a reception time difference (e.g., timing distributions) of thecoincidence photons detected by the first array of detectors (312) andsecond array of detectors (314) to classify the first location (334) orsecond location (336) as the origin of the positron annihilation event.The timing distributions of the first and second locations of theannular calibration source (330) are different from timing distributionsof positron emissions originating from a patient scan region (342) ofthe patient (340) because the annular calibration source (330) islocated closer to the detectors (312, 314) than the patient (340), andtherefore has a larger positron emission reception time difference thanthat of the emissions from the patient located closer to a center of thebore.

The first and second arrays of positron emission detectors may generatepositron emission data from the annular calibration source (330) and thepatient (340). A processor of the assembly (300) may distinguish thepositron emission data from the patient (340) and from the annularcalibration source (330). For example, the assembly (300) mayconcurrently classify (e.g., locate) the positron emission data from theannular calibration source (330) and the patient scan region (342)within the bore (322) using the reception times of the received photonpairs (e.g., using TOF data). A spatial filter may then be applied tothe calculated locations, as discussed in more detail below.

In some variations, the annular calibration source (330) may compriseone or more portions (e.g., an upper portion disposed above a plane ofthe support surface (350) and a lower portion disposed below the planeof the support surface (350)). In some of these variations, an upperportion of the annular calibration source (330) may be disposed abovethe patient scan region (342). For example, the annular calibrationsource (330) may be disposed at least about 20 cm above the supportsurface of the patient support (350) and/or above the highest point ofthe patient scan region (342) (e.g., above a plane intersecting thehighest point of the patient scan region (342) and parallel to thepatient support (350)). In another variation, a lower portion of theannular calibration source (330) may be disposed below a horizontalplane of the patient support (350). These exemplary annular calibrationsource (330) configurations may provide less additional dose to apatient relative to an annular calibration source (330) that encirclesthe patient (340) entirely. The annular calibration source (330) may beattached to the housing (320), and the housing (320) may be removablyattached to the gantry (310), thereby allowing the housing (320) andannular calibration source (330) to be replaced after its usefullifespan. The annular calibration source (330) may comprise a thicknessof about 0.10 mm to about 2.0 mm and a width of about 0.10 mm to about 5cm. In some variations, the annular calibration source (330) maycomprise a plurality of portions with each portion having differentshapes and dimensions.

C. Rotating Calibration Source

FIG. 4 is a cross-sectional schematic view of an assembly (400) that maybe configured to provide real-time calibration monitoring of PETdetectors and include one or more calibration source holders (408). Theradiation treatment assembly (400) may include a rotatable gantry (410)having mounted thereon a first array of positron emission detectors(412) and a second array of positron emission detectors (414) opposingthe first array of detectors (412). Each array of detectors (412, 414)may comprise a plurality of positron emission detectors. In somevariations, one or more calibration source holders (408) may be coupledto the gantry (410) such that the one or more calibration source holders(408) are fixed relative to the first array and the second array ofdetectors (412, 414). In other words, the positron emission detectors(412, 414) and calibration source holders (408) rotate together.

A housing (420) may couple to the gantry (410) (e.g., the housing (420)may be disposed over the gantry (210)), and the housing (220) maycomprise a bore (422) in which a patient (440) and a patient support(450) may be disposed. A patient scan region (442) (e.g., patientimaging field of view) may be represented by a volume of the patient(440) and/or patient support (450) detectable by the detectors (412,414). The patient support (450) may comprise a movable support surfaceon which the patient (440) may be disposed and the movable supportsurface may be coupled to a base (not shown).

As illustrated in FIG. 4, the one or more calibration source holders(408) and/or calibration sources (430) (e.g., radiation sources) may bespaced away from the patient scan region (442) within the bore (422) andwithin a field of view of the positron emission detectors (412, 414).For example, one or more calibration source holders (408) and/orcalibration sources (430) may be coupled to the detectors (412, 414) andmovable relative to the housing (420) (e.g., rotate about the housing(420)). For example, the one or more calibration source holders (408)and/or calibration sources (430) may be disposed along the detectors(412, 414) on a side facing an opposing detector (414, 412).

At least some of the lines of response (432) originating from one ormore calibration sources (430) extend toward the first array ofdetectors (412) and the second array of detectors (414). As depicted inFIG. 4, at least some of the calibration source LORs (432) arenon-intersecting with the patient scan region (442) and/or patientsupport (450). LORs (434) may be emitted from an ROI (444) of thepatient (440). LORs (432) emitted by the calibration sources (430) maybe detected by the positron emission detectors (412, 414). Consequently,each calibration source holder (408) is located for a calibration source(430) such that at least one LOR (432) emitted by the calibration source(430) is unobstructed by the patient (440) and/or the patient support(450). In some variations, one or more calibration source holders (408)may be located along a line non-intersecting with a patient imagingfield of view (e.g., patient scan region (442)) of the detectors (412,414), and the line may extend from the first array of detectors (412) tothe second array of detectors (414). The patient imaging field of viewmay comprise a volume located between the first array and second arrayof detectors. In other words, one or more calibration source holders(430) may be fixed relative to the first array and the second array ofdetectors (412, 414) and located with at least one line of sightunobstructed by a patient (440) and/or patient support (450), althoughthe line may pass through other structures such as the housing (420).

It should be noted that the assembly (400) may comprise a singlecalibration source holder (408) where the exemplary FIG. 4 illustratesfour calibration source holders (408) and calibration sources (430). Thecalibration source holder (408) may be configured to hold thecalibration source (430) in a manner as described in detail above. Thecalibration source holder (408) may allow the calibration source (430)to be removably coupled to the detectors (412, 414) such that thecalibration source (430) may be replaced after its useful lifespan.Although not illustrated in FIG. 4, the gantry (410) may furthercomprise a treatment radiation source and treatment detector between thefirst array and second array of detectors (412, 414) having a treatmentfield of view (460).

II. Methods

Also described here are methods for generating a fault signal duringdelivery of treatment radiation and/or patient imaging using the systemsand assemblies described herein. The system may comprise first andsecond arrays of PET detectors mounted on a rotatable gantry, where thePET detectors are configured to receive positron emissions from acalibration source (e.g., radiation source) and/or a patient to generatecalibration data and/or to extract positron emission data representingpositron emission activity originating from the calibration radiationsource. That is, calibration data may comprise data extracted from theacquired positron emission data that pertains to the calibration source,and in some variations, may also include any data quantity calculatedbased upon acquired positron emission data, such as calculated locationdata of the calibration source and/or patient. In some variations,positron emissions from the calibration source may be distinguished fromthe positron emissions from a patient by identifying the LORs that donot cross or intersect with a patient scan or treatment region of thebore. Examples of calibration data may include spatial resolution data,temporal resolution data, energy sensitivity data and/or energyprecision data, and/or the ability to accurately determine the locationof the calibration radiation source. In some variations, calibrationdata may include the number of detected LORs that intersect with thelocation of a stationary calibration source over one or more gantryrotations, and/or the energy level(s) of the detected LORs (i.e., thecoincident photon events) over one or more gantry rotations, and/or thenumber of detected LORs or coincident photon events detected withincoincidence time windows having different window widths. If thecalibration data exceeds a threshold parameter, such as when thecalibration data does not sufficiently correspond to a referencelocation or location range of the calibration source for a predeterminedtime period, a fault signal may be generated.

Other examples of calibration data deviations that may trigger thegeneration of a fault signal may include detecting a greater (or lesser)number of LORs that intersect with the location of the stationarycalibration source than is expected (based on the known and specifiedpositron emission rate and/or radioactivity of the calibration source),detecting LORs having energy level(s) that deviate from a 511 keV peak,and/or not detecting a threshold number of LORs or coincident photonevents with specific time difference values (e.g., time differences thatcorrespond to one or more of PET detectors being close to thecalibration source while other PET detectors are located far from thecalibration source) and/or within a first coincidence time windowcentered around a time difference value of 0 ns (i.e., the number of“true” coincident photon events). Alternatively or additionally, if thenumber of coincident photon events detected within a second coincidencetime window centered around a non-zero time different value (i.e., thenumber of “random” coincident photon events) exceeds the number ofcoincident photon events detected within the first coincidence timewindow (i.e., the number of “true” coincident photon events), a faultsignal may be generated. The system may respond in one or more ways tothe fault signal including performing additional diagnostics on thedetectors and/or assemblies (e.g., gantry motion sensor and/or encoderdiagnostics/characterization, calibration radiation sourcediagnostics/characterization, etc.), deactivating one or more detectors,stopping radiation therapy treatment, and/or recalibrating thedetectors. It should be appreciated that any of the systems describedherein may be used to determine a fault in the detectors using themethods discussed below.

FIGS. 5A-5B illustrate a fault detection method (500) using any of theassemblies (100, 200, 300, 400) discussed above. For example, aradiation treatment assembly may comprise a gantry comprising aplurality of positron emission detectors, an imaging radiation source, atreatment radiation source, and one or more calibration sources providedbetween the detectors. For example, one or more calibration sources maybe disposed in a housing within a field of view of the positron emissiondetectors and located at the same location along the length of theassembly as the positron emission detectors. The positron emissiondetectors and one or more calibration sources may be configured invarious combinations of stationary and rotatable elements, as discussedin detail above. Prior to an imaging and/or radiation therapy procedure,a patient may be administered a radioisotope for uptake into the body.The patient may lay on a patient support (e.g., couch) and be moved intoa bore of a gantry such that a ROI of the patient (e.g., one or moretumors or lesions) may be within a field of view of the positronemission detectors. Both the patient and the calibration source may beprovided between the detectors with the calibration source spaced awayfrom the patient. For example, the patient may be disposed within a boreof the assembly and the calibration source held by a calibration sourceholder on a surface of the housing.

An array of positron emission detectors may be activated to detectpositron emissions emitted within a detector field of view (502). Insome variations, the positron emission detectors may be mounted on arotatable gantry where the gantry is configured to rotate about a boreof the gantry while in other variations the detectors may be mounted ona stationary gantry.

The positron emission detectors may receive positron emission data fromany positron emission source provided between the detectors. Forexample, positron emission data from the patient and the calibrationsource may be continuously received and processed (504). The positronemission data may include TOF data such as a detector reception locationand reception time of the photon by that detector. Concurrently, thepatient may be treated by receiving a treatment radiation dose from thetreatment radiation source (506). That is to say, positron emission datamay be received during a treatment session. While positron emission datamay be acquired during a treatment session, the timing of the positronemission data acquisition may be staggered with respect to the timing ofthe treatment beam pulses. For example, positron emission data may bereceived by the positron emission detectors at predetermined timeintervals between treatment radiation beamlets. The patient may receiveradiation therapy treatment from treatment modalities including, forexample, emission-guided radiation therapy (EGRT), stereotactic bodyradiation therapy (SBRT), and intensity modulated radiation therapy(IMRT). Additionally or alternatively, the patient may be concurrentlyimaged by an imaging radiation source. The patient treatment and/orimaging step (506) may be performed in parallel (e.g., within apredetermined time period) with one or more of the steps of FIGS. 5A-5B.For example, the method (500) may perform one or more imaging andtreatment steps (506) with the reception of the positron emission data(504) and/or generation of calibration data (516) with at least sometemporal overlap. As another example, the step of receiving positronemission data from the calibration source and from the patient (504) maybe performed concurrently with generating a fault signal (524).

An annihilation event comprising a photon pair may be detected byopposing positron emission detectors and define an LOR. LORs of thepositron emission data may be classified by a processor to one or morecalculated locations or location ranges of the calibration source andthe patient (508). As discussed in more detail below, calculatedpositron emission locations may be classified by a processor usingreference time offsets or a spatial filter (509). In some variations,the LORs are classified using positron emission data received within acoincidence time window of between about 6 nanoseconds and 10nanoseconds.

Reference Time Offsets

The trajectory of a photon pair emitted by a positron annihilation eventdefines a LOR (e.g., a positron emission path). The origin of theannihilation event is at point located along the LOR and may be locatedusing a difference in photon reception times (i.e., a time difference ortime offset). In particular, a difference in reception times between thephotons corresponds to a distance traveled by the photons along the LORfrom the annihilation event. For example, if each photon of anannihilation event arrives at a respective detector at the same time,then the photons have traveled an equal distance such that theannihilation event originates at a midpoint of the LOR. Accordingly, insome variations, a positron emission (e.g., annihilation event) originlocation corresponding to the calibration source and/or the patient maybe calculated by a processor using TOF data (e.g., reception timedifference of the pair of photons), LOR, and reference time offsets(510). For example, positron emission path data may be classified asoriginating from a calibration source (e.g., stationary radiationsource) using a difference between a reception time of the pair ofphotons with a time threshold parameter range (e.g., time differencerange) corresponding to the calibration source. Positron emission pathdata outside the time threshold parameter range may be classified asnoise or as originating from the patient. In some variations, thepositron emission detector may comprise a coincidence time resolution ofat least about 300 picoseconds (ps) (FWHM) corresponding to a positionaluncertainty of about 4.5 cm (FWHM) along the LOR.

In some of these variations, the calibration source may be distinguishedfrom the positron emission data by comparing the difference in receptiontimes (e.g., time difference or time offset) of a photon pair to areference time offset corresponding to a location of the calibrationsource. For example, a look-up table (LUT) of reference time offsets(calculated using the equations discussed in detail below) may be storedin memory and compared to positron emission data. In some variations, astationary calibration source may be located on a housing of a rotatablegantry relative to the PET detectors rotating about the patient scan ortreatment region such that the time offset (e.g., time difference) isabout 2.5 ns when PET detectors are located at specified gantry anglesabout the patient scan or treatment region. It should be understood thatin other variations with different arrangements of PET detectors and/oremission properties calibration source(s), the time difference or timeoffset value corresponding to coincident photon originating from thecalibration source may have different time difference values. In somevariations, if the number of coincident photon events detected for timeoffsets other than 2.5 ns (“random” coincident photon events) is thesame as, or greater than, the number of coincident photon eventsdetected for a time offset of 2.5 ns (“true” coincident photon events),a fault signal may be generated. In some variations, the coincidentphoton events originating from the calibration source (e.g., the peakscentered around a time difference value of 2.5 ns in a histogram thatplots the coincidence photon event count over a coincidence time window)may be distinguished from the coincident photon events originating froma PET-avid region in a patient using the spatial/sinogram filtersdescribed herein, optionally in combination with temporal filters thatselect for events with particular temporal characteristics (e.g.,certain time difference values between the detection times of the twophotons in a coincident event). In some variations, a passing criteriaor threshold may be a ratio of the number of true coincident photonevents to the number of random coincident photon events (true-to-randomratio value), where the value of the ratio is greater than or equal toone. It should be understood that while the examples described hereinpertain to a system where the positron emission events detected by tworotatable PET detector arrays that originate from a single stationarycalibration source have a time difference or offset value of about 2.5ns, it should be understood that in other systems with different sizesand/or relative positions between the PET detector arrays and one ormore calibration sources, the time difference or offset value may be anyvalue (e.g., any value other than 2.5 ns). The width of the coincidencetime window for detecting coincident photon events from either or boththe calibration source(s) and/or PET-avid patient or phantom within thebore may be adjusted (e.g., expanded or narrowed relative to theexamples described herein) as appropriate so that these coincidentphoton events may be extracted from the PET detector data. While thecoincidence time window width may vary, the coincidence time windowwidth may be centered around about 0 ns.

A. Stationary Calibration Source

In variations where a radiation treatment assembly comprises a rotatablegantry having mounted thereon positron emission detectors and astationary housing coupled to a calibration source (e.g., FIGS. 2A-2B),a distance of the detector to the calibration source may vary as afunction of gantry angle as the PET detectors rotate about a bore of thegantry. From a positron emission detector point-of-view (e.g., PETdetector reference frame), the position of the PET detector is fixed asthe calibration source rotates at a gantry angle theta (θ) definedrelative to a reference direction. With respect to the PET detectorreference frame, the location of the detector may be defined bycoordinates <x,y> where the X-axis corresponds to a horizontal axisplane (e.g., parallel to the ground and/or a patient support) and theY-axis corresponds to a vertical plane perpendicular to the horizontalplane. The location of the calibration source may be defined as

r cos θ, r sin θ

where r represents a radius from a reference point (e.g., origin of thedetector <x,y>) and θ is an angle from the reference direction. Alocation of a photon pair (e.g., first and second photons) detected byrespective detectors may be defined as <x₁,y₁> and <x₂,y₂>,respectively. A distance from the first and second photon detectionlocation to the calibration source is given by:

d ₁=√{square root over ((r cos θ−x ₁)²+(r sin θ−y ₁)²)}

d ₂=√{square root over ((cos θ−x ₂)²+(r sin θ−y ₂)²)}

The first and second photons each travel at the speed of light (c) fromthe annihilation point to opposing positron emission detectors. Adifference in reception time (i.e., a time difference or time offset)between the first and second photons is given by:

Δt=2*(d ₁ −d ₂)/c

Accordingly, a table of reference time offsets Δt may be calculated as afunction of the gantry angle θ and photon pair distances d₁, d₂:

t(θ,d ₁ ,d ₂)=Δt

A reference time offset may be compared to a corresponding detectedphoton pair time offset and a threshold parameter to generate a faultsignal, as described in more detail below.

B. Annular Calibration Source

In other variations, an assembly may comprise a rotatable gantry havingmounted thereon positron emission detectors and a stationary housingcoupled to a calibration source such as the annular radiation sourceillustrated in FIG. 3. Each LOR of the stationary calibration source mayoriginate from one of two locations (e.g., where the LOR intersects theannular radiation source). Therefore, the LUT includes two possiblereference time offsets (e.g., locations) for each gantry angle θ andphoton pair distances d₁, d₂ rather than a single reference time offset.

t(θ,d ₁ ,d ₂)=Δt ₁ ,Δt ₂

For a stationary calibration source, a LUT stored in memory may comprisea pair of reference time offsets (Δt₁, Δt₂) for each set of gantry angleand photon pair distances. Due to the wide spatial and temporalseparation between the two calibration source intersection points, apositron emission origin may be compared to the closest referencecalibration source location with a small probability of error.

C. Rotating Calibration Source

In other variations, an assembly may comprise a rotatable gantry havingmounted thereon both positron emission detectors and one or morecalibration sources (e.g., radiation source, radiation source holder),such as illustrated in FIG. 4, where the calibration source is fixedrelative to the positron emission detector. In this configuration, thedistance of the calibration source to the detector does not vary (from aPET detector reference frame) and therefore does not require the LUT andcalculations using the equations described above.

Patient Spatial Filter

In some variations, a positron emission origin location corresponding tothe calibration source and/or the patient may be calculated using TOFdata and one or more spatial filters. In other words, the calculatedlocation of an annihilation event based on TOF data may be compared toreference locations or location ranges of the patient and calibrationsource and be used to classify the positron emissions and determinewhether the coincident event originated from the patient or calibrationsource. In some variations, given the locations of a set of positronemissions, one or more spatial filters or classifiers may be applied todistinguish the positron emission data from the calibration source(e.g., radiation source and/or radiation source holder) and the patient(e.g., patient scan region). For example, a first spatial filter of thecalibration source and a second spatial filter of the patient may beapplied to distinguish the sources of the positron emission data. Insome of these variations, positron emissions outside of the first andsecond spatial filters may be excluded as noise. The spatial filtersmay, for example, comprise a volume of the calibration source and/orpatient and include a margin for error (predetermined or adjustable). Insome variations, the margin may be from about 1 cm to about 10 cm.Alternatively or additionally, a spatial filter may designate a regionwithin a bore of the system where the calibration source is expected tobe located, and the processor may extract the LORs that intersect thatregion of the spatial filter but do not intersect the portion of thefilter corresponding to the patient scan or treatment region,designating the extracted LORs as containing data pertaining thepositron emission activity originating from the calibration source. Insome variations, the processor may count the number of LORs thatintersect that region of the spatial filter, and if that count does notmeet or exceed a threshold of expected LORs (e.g., calculated based onthe known radioactive and/or positron emission properties of thecalibration source), a fault signal may be generated.

In some variations, the second spatial filter of the patient maycomprise a geometry using a patient treatment plan including geometricand dosimetric parameters for radiation therapy treatment. For example,the patient treatment plan may comprise a set of target volumes (e.g., apatient region), including (in order of descending volume) an irradiatedvolume, treated volume, planning treatment volume (PTV), clinical targetvolume (CTV), and gross tumor volume (GTV). The PTV encompasses the GTVand adds a margin for error including patient alignment, machinecalibration, and a motion envelope of the ROI. Positron emissions fromthe GTV may be used to direct radiation beamlets in EGRT while positronemissions outside of the GTV and otherwise within a volume of the PTVmay be classified as noise. Position emissions located outside the PTVmay be rejected altogether.

In some variations, the geometry of the spatial filter may be adjustedby a processor using the patient treatment plan (512). For example, thesecond spatial filter of the patient may automatically adjust thegeometry of the second spatial filter to correspond to one of the GTV,CTV, and PTV, and may automatically adjust with modifications to thepatient treatment plan. Additionally or alternatively, the spatialfilter may be user adjustable such that a user such as an operator mayinput the volume of one or more of the first and second spatial filter.

In some variations, a processor may concurrently distinguish positronemission data from the calibration source and the patient using one ormore spatial filters (514). For example, one or more of the first andsecond spatial filters may be applied to the positron emission data toclassify the positron emission locations as corresponding to thecalibration source and/or patient. Positron emission data locatedoutside of a spatial filter (e.g., located outside a calibration regionand/or patient region) may be excluded as noise. In some of thesevariations, a first spatial filter may be applied to the positronemission data and all other positron emissions (such as from thepatient) may be excluded and disregarded.

Regardless of whether the positron emissions are classified usingreference time offsets (510) or spatial filters (514), the positronemission data from the calibration source may be used by a processor togenerate calibration data (e.g., a set of positron emission locations ofthe calibration source (516)). Optionally, the positron emission datafrom the patient may be used by a processor to generate patient data(e.g., a set of positron emission locations of the patient (518)). Thepatient data may be used to treat the patient (520) using, for example,a treatment radiation source under an EGRT procedure. It should be notedthat steps (518, 520) are optional for PET detector calibrationmonitoring.

A calibrated positron emission detector should classify positronemissions from a calibration source at a location that matches the truelocation of the calibration source within a predetermined margin oferror. A fault signal may be generated when a calculated location of thecalibration source does not correspond to a reference location orlocation range of the calibration source. In particular, a fault signalmay be generated (524) by a processor when the positron emissiondetector (e.g., calibration data) from the calibration source (e.g.,radiation source, radiation source holder) exceeds a threshold parameter(522—Yes). In some variations, the threshold parameter may be avariability threshold parameter. For example, the variability thresholdparameter may comprise a percentage of time that a calculated locationof the calibration source is outside a reference location or locationrange (e.g., volume) of the calibration source (e.g., proportion of timewhen the calculated and reference locations or location ranges of thecalibration source do not match). In some of these variations, a marginmay be added to the reference location or location range. In somevariations, a variability threshold may be about 0.1%, about 0.5%, about1.0%, about 1.5%, about 2%, about 5%, and values in between. In somevariations, calibration may be monitored continuously in time windows ofabout 1 seconds, about 10 seconds, about 10 seconds, and values inbetween. The variability threshold parameter may comprise one or more ofa time threshold and/or location deviation threshold. Therefore, thefault signal may be generated when the positron emission data exceedsone or more of a time threshold and location deviation threshold. Forexample, a fault signal may be generated when a mean reception timeoffset (Δt) of positron emissions over a predetermined time windowexceeds a time offset threshold parameter corresponding to a referencetime offset value.

In some variations, calibration data (for a single stationarycalibration source in a system with a rotatable gantry with two arraysof PET detectors, for example) may comprise the number of coincidentphoton events detected by one or more PET detectors over a specifiedtime period and/or gantry rotation. The expected number of coincidentphoton events to be detected for each PET detector (e.g., each detectormodule or detector crystal) from the extracted LORs (i.e., LORs thatintersect with the known location of the stationary calibration source)may be calculated based on the known properties of the calibrationsource as well as its location relative to the rotating PET detectors.The threshold parameter may be a tolerance or difference value betweenthe detected number of LORs that intersect with the calibration sourcelocation and the expected number of coincident photons emitted from thecalibration source. If the number of LORs detected by the PET detectorsdeviate from the expected number of expected emitted coincident photonsby more than the specified tolerance or difference value, then a faultsignal may be generated. For example, if the number of detected LORs pergantry angle position deviate from what is expected by more than thetolerance or difference threshold, and/or if the number of detected LORsover one or more gantry rotations deviate from what is expected by morethan the tolerance or difference threshold, then a fault signal may begenerated.

In some variations, alternatively or additionally, calibration data (fora single stationary calibration source in a system with a rotatablegantry with two arrays of PET detectors, for example) may comprise thenumber of coincident photons at particular energy levels (e.g., 511 keV)detected by one or more PET detectors over a specified time periodand/or gantry rotation. The expected number of coincident photon eventsto be detected for each PET detector (e.g., each detector module ordetector crystal) at various energy levels from the extracted LORs(i.e., LORs that intersect with the known location of the stationarycalibration source) may be calculated based on the known properties ofthe calibration source as well as its location relative to the rotatingPET detectors. The threshold parameter may be a tolerance or differencevalue between the detected number of coincident photons having an energylevel of 511 keV and the expected number of coincident photons emittedfrom the calibration source having an energy level of 511 keV. If thenumber of coincident photons having an energy level of 511 keV detectedby the PET detectors deviate from the expected number of expectedemitted coincident photons having an energy of 511 keV by more than thespecified tolerance or difference value, then a fault signal may begenerated. For example, if the number of detected coincident photonshaving an energy level of 511 keV per gantry angle position deviate fromwhat is expected by more than the tolerance or difference threshold,and/or if the number of detected coincident photons having an energylevel of 511 keV over one or more gantry rotations deviate from what isexpected by more than the tolerance or difference threshold, then afault signal may be generated. For example, if no 511 keV coincidentphotons from LORs that intersect the location of the stationarycalibration source are detected, a fault signal may be generated.

In some variations, alternatively or additionally, calibration data (fora single stationary calibration source in a system with a rotatablegantry with two arrays of PET detectors, for example) may comprise thenumber of coincident photon events detected that have specific timedifference or offset values (i.e., the time difference between thedetection of each of the photons in a coincident pair). The calibrationdata may include the number of coincident photon events detected over arange of time difference or offset values. This range of time differenceor offset values may be referred to as a coincidence time window. Insome variations, the number of coincident photon events may be measuredover coincidence time windows having different ranges and/or centeredover various time difference values, and/or over one or more gantryrotations. A histogram of coincident photon events measured in acoincidence time window may be generated by detecting coincident photonevents using the PET detectors, comparing the detection time of each ofthe photons of a coincident pair to obtain a time difference value foreach coincident photon event, and plotting (e.g., binning) the number ofcoincident photon events for each time difference value within thecoincidence time window. Based on the known properties of thecalibration source as well as its location relative to the rotating PETdetectors, the number of coincident photon events having particular timedifference values that correspond with the relative distances betweenthe calibration source and each of the two arrays of PET detectors maybe greater than the number of coincident photon events having other timedifference values. For example, in a variation of a system where astationary calibration source is located at the top or bottom of a bore(i.e., at a substantial distance away from the center of the bore and/orsystem isocenter) and the two PET detector arrays rotate and moverelative to the stationary calibration source, the coincident photonevents that originate from the calibration source may have a timedifference value of, for example, 2.5 ns. In contrast, coincident photonevents originating from a PET-avid region in a patient (i.e., close tothe center of the bore and/or system isocenter) may have a timedifference close to about 0 ns, since both annihilation photons willtravel a similar distance before being detected by the first and secondPET detector arrays. In some variations, the time difference ofcoincident photon events originating from a PET-avid region in a patient(or phantom) may be from about 0.1 ns to about 1.5 ns, e.g., about 0.25ns, about 0.5 ns, about 0.75 ns, about 0.6 ns, about 1 ns, etc.,depending on the geometry and location of the patient within thetreatment and/or scanning region in the bore. FIG. 8A depicts oneexample of a system with a single calibration source at a bottom or top(i.e., edge region) of the bore, and two rotating PET detector arrayswithout a PET-avid region within the center of the bore, while FIG. 8Bdepicts an example of a system similar to that of FIG. 8A, but with aPET-avid region within the center of the bore (e.g., a PET-avid phantomor patient), resulting in a peak centered around the time differencevalue 0 ns.

In some variations, the coincident photon events detected over a rangeof time differences centered around the time difference value of 0 nsmay be classified as “true” coincident photon events, while thecoincident photon events detected over a range of time differencescentered around a non-zero time difference value may be classified as“random” coincident photon events. For example, if the relativepositions of the PET detector arrays and the stationary calibrationsource give rise to coincident photon events with a time difference of2.5 ns, then “true” coincident photon events may be events detectedwithin a coincidence time window that includes time difference valuesbetween −2.5 ns and +2.5 ns (centered around a time difference value of0 ns). Alternatively or additionally, a coincidence time windowcomprising “true” coincident photon events may have a range of timedifference values from −5 ns to +5 ns. Random coincident photon eventsmay be events detected within a coincidence time window between 15 nsand 25 ns (centered around a time difference value of 20 ns), or anytime window that is centered around a non-zero time difference value anddoes not overlap with the “true” coincidence time window. If the numberof coincident photon events in the “random” coincidence time window isthe same as, or greater than, the number of coincident photon events inthe “true” coincidence time window, a fault signal may be generated. Insome variations, a threshold criterion may be a ratio of the number ofcoincident photon events in the “true” coincidence time window (e.g., afirst coincidence time window centered around a 0 ns time difference oroffset) to the number of coincident photon events in the “random”coincidence time window (e.g., a second coincidence time window centeredaround a non-zero photon reception time difference or offset with arange that does not overlap with the first coincidence time window).This ratio may be referred to as a true-to-random ratio value, where apassing threshold criterion may be a ratio value that is greater than orequal to one. If the true-to-random ratio value does not meet or exceedone, then a fault signal may be generated. The true-to-random ratiovalue may be calculated by counting the number of coincident photonevents within the first coincidence time window (e.g., number of trueevents), counting the number of coincident photon events within thesecond coincidence time window (e.g., number of random events), anddividing the number of true events by the number of random events. FIGS.8A-8C depict plots (e.g., histograms) that represent the number (counts)of coincident photon events detected for various coincidence time windowwidths. FIG. 8A depicts one example of a histogram or plot of the numberof true coincident photon events detected within a first (true)coincidence time window have a range between −5 ns and 5 ns for astationary calibration source (no PET-avid patient) over one or moregantry rotations. As may be seen there, there are two peaks within thecoincidence time window, each of the peaks centering around +/−2.5 ns,respectively. The first peak (800) centered around −2.5 ns representsthe number of coincident photon events detected at a first gantrylocation of the first array of PET detectors where the first array ofPET detectors is located at its closest distance to the calibrationsource (i.e., while the second array of PET detectors is located at asignificantly greater distance away from the calibration source). Thesecond peak (802) centered around +2.5 ns represents the number ofcoincident photon events detected at a second gantry location of thesecond array of PET detectors where the second array of PET detectors islocated at its closest distance to the calibration source (i.e., whilethe first array of PET detectors is located at a significantly greaterdistance away from the calibration source).

FIGS. 8B and 8C depict similar plots, but in addition to the calibrationsource, also include a patient with a PET-avid region (e.g., a tumorregion that has elevated levels of PET activity due to preferentialuptake of a PET tracer). The peak (810) centered around the timedifference value of about 0 ns represents the coincident photon eventsdetected that originate from the patient. In these examples, the first(true) coincidence time window may be between −5 ns and 5 ns, centeredaround 0 ns. The second (random) coincidence time window may be between15 ns and 25 ns, centered around 20 ns. FIG. 8B depicts one variation ofa system where the PET detectors are functioning properly and/orcalibrated as desired. As may be seen there, within this firstcoincidence time window, the two peaks (812, 814) centered around −/+2.5ns represent the coincident photon events originating from thecalibration source and detected by the PET detectors. The number of“random” coincident photon events (816) in the second coincidence timewindow (e.g., between 15 ns and 25 ns, centered around 20 ns) is lessthan the number of coincident photon events in the first coincidencetime window; that is, the true-to-random ratio is about one or more.FIG. 8C depicts one variation of a system where the PET detectors arenot functioning properly and/or are no longer calibrated as desired. Asmay be seen there, there are no peaks centered around −/+2.5 ns in thefirst (true) coincidence time window (between −5 ns to 5 ns), and thenumber of “random” coincident photon events (818) in the second (random)coincidence time window (between 15 ns and 25 ns, centered around 20 ns)is greater than or equal to the number of coincident photon eventshaving a time difference of about 2.5 ns; that is, the true-to-randomratio is less than or equal to one. Since the peaks are notidentifiable, and/or are below a threshold value (i.e., less than thenumber of random coincident photon events), a fault signal may begenerated. In some variations, a threshold criterion for a passingtrue-to-random ratio may be a ratio value that is greater than one(e.g., about 1.1 or more, about 1.2 or more, about 1.3 or more, about1.5 or more, etc.) and a fault signal may be generated if theactual/calculated true-to-random ratio is less than that ratio value(e.g., less than about 1.1, less than about 1.2, less than about 1.3,less than about 1.5, etc., respectively). For example, if the thresholdratio value is 1.2, a fault signal may be generated if the actualtrue-to-random ratio is calculated to be 1.1 or 1. The generation of afault signal may include generating a notification to the user to checkthe calibration and/or proper functioning of the PET detectors, and/organtry motion sensors, and/or to check or further characterize the PETtracer uptake by the patient.

The calibration monitoring of the PET detectors disclosed herein isbased on a direct measurement of the detectors and may optionally beused to corroborate and/or verify other sensors in the radiation therapyassembly. For example, the radiation therapy assembly may comprise oneor more temperature sensors that monitor the temperature of the PETdetectors. In systems without a positron-emitting calibration source,these temperature readings may be used as an indicator of PET detectorfunction. For example, the temperature of the positron emissiondetectors may be measured using one or more temperature sensors, and ifthe temperature exceeds a threshold value (e.g., an increase of 2degrees during a procedure), a fault signal may be generated. However,monitoring PET detector function by measuring only secondary factors(such as temperature) may result in an inaccurate assessment if any ofthe sensors fail (especially due to the high-radiation environment inthe vicinity of a radiation therapy assembly). That is, a fault signalbased on temperature measurements may be the result of a fault in one ormore positron emission detectors and/or a fault in a temperature sensor.Calibration monitoring of the PET detectors using a positron-emittingcalibration source as described herein allows for a separateverification of a fault signal generated by the temperature sensor(s).This may provide two independent checks on positron emission detectorcalibration and increase confidence in calibration monitoring.Alternatively or additionally, a radiation therapy assembly may compriseone or more sensors that monitor parameters and various subsystem of thetherapy assembly. Some radiation therapy assemblies may monitor thetemperature of other portions of the radiation therapy assembly (e.g.,gantry, imaging source and detector, multi-leaf collimator, bore volume,regions in the vicinity of the PET detectors, etc.), the flowrate of acooling fluid (using one or more flow meters), and/or ambient and/orscattered radiation levels (using one or more radiation detectors suchas scintillation counters, Gieger counters, gaseous ionizationdetectors, ionization chambers, and the like). Data readings from one ormore of the sensors may be used in conjunction with calibration sourceemission data to assess and/or adjust one or more of the PET detectors.A fault signal generated based on signals detected by the PET detectorsmay also trigger a user to check the calibration of the motion and/orposition sensors and/or position encoders of a rotatable gantry.

A radiation treatment assembly may respond in one or more ways inresponse to the generation of a fault signal (524). For example, theradiation treatment assembly may deactivate one or more positronemission detectors, output the detector status to an operator, stop animaging and/or radiation therapy treatment procedure, and calibrate thedetector using the calibration data. In some variations, an interlockmay be generated and visually represented on a display to the user. Forexample, a visual or graphical user interface depicted on a display ormonitor may comprise a status bar or icon of each subsystem of theradiation treatment assembly. The appearance of the status bar or iconmay be updated at regular time intervals and/or as desired by the user.When a fault signal is generated and an interlock is triggered, thestatus bar or icon for that subsystem may change color (e.g., turn red).The user may then click on the status bar or icon to obtain adescription of the error and/or to commence further testing and/orcalibration of that subcomponent. When a fault is detected using one ormore of the methods described herein, the icon for the PET detectors(e.g., one icon per PET detector array), and/or the icon for the gantrymotion/position encoders, and/or the icon for the calibration source,and any other subsystem may change color or form. The user may click onany one of the icons to troubleshoot or obtain further details about thestatus of that subsystem. Alternatively or additionally, a fault signalmay also prompt the user to confirm that the PET tracer was properlyintroduced to the patient.

In some variations, the fault signal may comprise one or more positronemission detectors at fault. The number of faulty detectors may becompared to a predetermined threshold (526) to determine which faultydetectors to deactivate. In some of these variations, at least one ofthe positron emission detectors of a first and second array of thedetectors may be deactivated based on the fault signal (528, 530). Inone example, up to three positron emission detectors may be deactivatedbased on the fault signal indicating fault in up to three of thepositron emission detectors (528). Conversely, each of the first arrayand second array of detectors may be deactivated based on the faultsignal indicating fault in four or more of the detectors (530).Deactivating a subset of the positron emission detectors allows patientimaging and/or radiation therapy treatment to continue when some of thepositron emission detectors are uncalibrated. It should be appreciatedthat the threshold number of faulty detectors may be any number ofdetectors. For example, if even a single PET detector is faulty oruncalibrated, the delivery of treatment radiation may be paused and/orthe treatment session may be stopped.

In some variations, radiation therapy treatment of the patient using thetreatment radiation source may be stopped (532) in response to the faultsignal. For example, radiation therapy treatment may be stopped withinabout 0.01 seconds after the fault signal is generated, therebysignificantly reducing potential harm to the patient due to loss ofcalibration of one or more positron emission detectors. Conventionalquality assurance (QA) procedures that verify calibration once a day areunable to monitor PET detector calibration in real-time.

In some variations, the radiation treatment assembly may undergo a QAprocedure to verify the fault signal and then recalibrate the faultydetectors if necessary. One or more positron emission detectors may becalibrated by a processor using the calibration data (534). For example,a processor may recalibrate one or more of the positron emissiondetectors (e.g., dynamic range of light detectors) by compensating for aconsistent time offset error in the calibration data. Other aspects ofthe one or more positron emission detectors that may be adjusted orcorrected may include energy calibration, energy resolution, detectorcount-rate uniformity, and/or dead-time corrections.

In some variations, a detector status may be output (536) based on oneor more of the fault signal, positron emission detector deactivation,radiation therapy treatment status, and positron emission detectorcalibration. One or more visual, audio, and tactile sensory outputsystems coupled to the radiation treatment assembly may be used tooutput the detector status to a user such as an operator. For example, adisplay coupled to the radiation treatment assembly may display thedetector status to an operator. A detector status may be outputcontinuously, at predetermined intervals, upon a change in status, andupon generating a fault signal. Additionally or alternatively, thedetector status may be stored in memory and/or transmitted over anetwork to be output and/or displayed to one or more of a remoteoperator, system vendor, regulatory agency, and/or stored in a database.

Some variations described herein relate to a computer storage productwith a non-transitory computer-readable medium (also may be referred toas a non-transitory processor-readable medium) having instructions orcomputer code thereon for performing various computer-implementedoperations. The computer-readable medium (or processor-readable medium)is non-transitory in the sense that it does not include transitorypropagating signals per se (e.g., a propagating electromagnetic wavecarrying information on a transmission medium such as space or a cable).The media and computer code (also may be referred to as code oralgorithm) may be those designed and constructed for the specificpurpose or purposes. Examples of non-transitory computer-readable mediainclude, but are not limited to, magnetic storage media such as harddisks, floppy disks, and magnetic tape; optical storage media such asCompact Disc/Digital Video Discs (CD/DVDs), Compact Disc-Read OnlyMemories (CD-ROMs), and holographic devices; magneto-optical storagemedia such as optical discs; solid state storage devices such as a solidstate drive (SSD) and a solid state hybrid drive (SSHD); carrier wavesignal processing modules; and hardware devices that are speciallyconfigured to store and execute program code, such asApplication-Specific Integrated Circuits (ASICs), Programmable LogicDevices (PLDs), Read-Only Memory (ROM) and Random-Access Memory (RAM)devices. Other variations described herein relate to a computer programproduct, which may include, for example, the instructions and/orcomputer code disclosed herein.

The systems, devices, and/or methods described herein may be performedby software (executed on hardware), hardware, or a combination thereof.Hardware modules may include, for example, a general-purpose processor(or microprocessor or microcontroller or multi-core processor), a fieldprogrammable gate array (FPGA), and/or an application specificintegrated circuit (ASIC). Software modules (executed on hardware) maybe expressed in a variety of software languages (e.g., computer code),including C, C++, Java®, Ruby, Visual Basic®, and/or otherobject-oriented, procedural, or other programming language anddevelopment tools. Examples of computer code include, but are notlimited to, micro-code or micro-instructions, machine instructions, suchas produced by a compiler, code used to produce a web service, and filescontaining higher-level instructions that are executed by a computerusing an interpreter. Additional examples of computer code include, butare not limited to, control signals, encrypted code, and compressedcode.

Although the foregoing variations have, for the purposes of clarity andunderstanding, been described in some detail by of illustration andexample, it will be apparent that certain changes and modifications maybe practiced, and are intended to fall within the scope of the appendedclaims. Additionally, it should be understood that the components andcharacteristics of the systems and devices described herein may be usedin any combination. The description of certain elements orcharacteristics with respect to a specific figure are not intended to belimiting or nor should they be interpreted to suggest that the elementcannot be used in combination with any of the other described elements.For all of the variations described above, the steps of the methods maynot be performed sequentially. Some steps are optional such that everystep of the methods may not be performed.

1. An radiotherapy system comprising: a rotatable gantry; a first arrayof positron emission detectors mounted on the gantry and a second arrayof positron emission detectors mounted on the gantry opposite the firstarray of positron emission detectors; a therapeutic radiation sourcemounted on the rotatable gantry between the first and second arrays ofpositron emission detectors; a housing disposed over the rotatablegantry and comprising a bore and a stationary radiation source holderspaced away from a patient region within the bore, wherein thestationary radiation source holder is located within the housing or on asurface of the housing; and a processor configured to receive positronemission data detected from the first and second arrays of positronemission detectors and to extract positron emission data representingpositron emission activity originating from the stationary radiationsource holder, and to generate a fault signal when the extractedpositron emission data does not satisfy one or more threshold criteria.2. The system of claim 1, further comprising a patient support, thepatient support comprising a movable support surface and a base.
 3. Thesystem of claim 2, wherein the radiation source holder is disposed alongthe surface of the housing at a location above the patient scan region.4. The system of claim 2, wherein the radiation source holder is locatedbelow the movable support surface.
 5. The system of claim 1, furthercomprising a calibration radiation source held by the radiation sourceholder, the calibration source comprising a radioactivity of about 1 μCito 300 μCi.
 6. The system of claim 1, further comprising a calibrationradiation source configured to be retained by the radiation sourceholder, the calibration source comprising a shape with a maximumdimension from about 0.25 inch to about 3 inches.
 7. The system of claim6, wherein the calibration radiation source comprises a disk-shapedenclosure and a positron-emitting element located within the enclosure.8. The system of claim 1, wherein the processor is further configured toconcurrently extract the positron emission data representing positronemission activity originating from the radiation source holder and toextract positron emission data representing positron emission activityoriginating from the patient scan region.
 9. The system of claim 1,wherein a threshold criterion comprises a spatial filter that selectsfor positron emission activity originating from a location of thestationary radiation source holder, and wherein a fault signal isgenerated when applying the spatial filter to the extracted positronemission data indicates that the positron emission activity does notco-localize with the location of the stationary radiation source holder.10. The system of claim 9, wherein the spatial filter is useradjustable.
 11. The system of claim 9, wherein the processor is furtherconfigured to automatically adjust a geometry of the spatial filterusing a patient treatment plan.
 12. The system of claim 1, wherein thefirst and second arrays of positron emission detectors define an imagingplane, wherein a beam of the therapeutic radiation source defines atreatment plane, and the imaging plane and the treatment plane areco-planar.
 13. The system of claim 12, wherein the stationary radiationsource holder is co-planar with the imaging plane and the treatmentplane.
 14. The system of claims 5-7, wherein the stationary radiationsource holder comprises a groove having a shape that corresponds with ashape of the radiation source.
 15. The system of claim 1, wherein athreshold criterion comprises a threshold number of coincident photonevents detected with a first time difference, wherein the processor isconfigured to generate a plot of an actual number of coincident photonevents detected with the time difference, and wherein a fault signal isgenerated when the actual number of coincident photon events occurringwith the time difference does not exceed the threshold number.
 16. Thesystem of claim 15, wherein a threshold criterion comprises a thresholdtrue-to-random ratio value, wherein the processor is configured togenerate a ratio of the actual number of coincident photon eventsoccurring within a first coincidence time window centered around 0 ns toan actual number of coincident photon events occurring within a secondcoincidence time window that does not overlap with the first coincidencetime window, and wherein a fault signal is generated if the ratio doesnot exceed the threshold true-to-random ratio value.
 17. The system ofclaim 16, wherein the threshold true-to-random ratio value is about 1.18. The system of claim 1, wherein a threshold criterion comprises afirst expected number of coincident photon events to be detected with afirst detection time difference of about 2.5 ns at a first gantrylocation of the first array of positron emission detectors and a secondexpected number of coincident photon events to be detected with adetection time difference of about 2.5 ns at a second gantry location ofthe first array of the positron emission detectors that is 180° from thefirst gantry location, wherein the processor is configured to generate aplot of actual numbers of coincident photon events detected within acoincidence time window between −5 ns to +5 ns over a 360° gantryrotation based on positron emission data detected by the first andsecond arrays of positron emission detectors, and wherein a fault signalis generated when an actual number of coincident photon events detectedwith a detection time difference of about 2.5 ns at the first gantrylocation of the first array of the positron emission detectors does notmeet or exceed the first expected number, and an actual number ofcoincident photon events detected with a detection time difference ofabout 2.5 ns at the second gantry location of the first array of thepositron emission detectors does not meet or exceed the second expectednumber.
 19. The system of claim 1, wherein a threshold criterioncomprises an expected number of coincident photon events to be detectedby each positron emission detector of the first and second arrays ateach gantry location over a 360° gantry rotation, wherein the processoris configured to calculate, using the positron emission data detected bythe first and second array of positron emission detectors, an actualnumber of coincident photon events detected by each positron emissiondetector of the first and second arrays at each gantry location over a360° gantry rotation, and wherein a fault signal is generated when adifference between the actual number of coincident photon events and theexpected number of coincident photon events exceeds a predetermineddifference threshold for at least one positron emission detector. 20.The system of claim 1, wherein a fault signal is generated when theprocessor does not detect any positron emission data representingpositron emission activity originating from the stationary radiationsource holder.
 21. The system of claim 1, wherein a threshold criterioncomprises an energy resolution spectrum with a coincident 511 keV photonevent count above a peak threshold, and wherein a fault signal isgenerated when an energy resolution spectrum generated from the positronemission data does not have a 511 keV photon event count above the peakthreshold.
 22. The system of claim 1, further comprising a display andwherein the processor is configured to generate a visual indicator andtransmitting the visual indicator to the display, wherein the visualindicator has a first appearance in the absence of a fault signal and asecond appearance different from the first appearance when a faultsignal is generated.
 23. An imaging assembly comprising: a gantrycomprising a first array of rotatable positron emission detectors and asecond array of rotatable positron emission detectors opposing the firstarray of detectors; a housing disposed over the gantry and comprising abore and a stationary radiation source spaced away from a patient scanregion within the bore, wherein the stationary radiation source islocated within the housing or on a surface of the housing; and aprocessor configured to receive positron emission path data from thefirst and second arrays of rotatable positron emission detectors and toclassify positron emission path data that originates from the stationaryradiation source, and to generate a fault signal when the stationaryradiation source positron emission path data exceeds a thresholdparameter.
 24. The assembly of claim 23, wherein a pair of photonsemitted by a positron annihilation event generates a positron emissionpath, and wherein the processor is configured to classify the positronemission path data that originates from the stationary radiation sourceusing a difference between a reception time of the pairs of photonswithin a time threshold parameter range.
 25. The assembly of claim 24,wherein the threshold parameter is a location deviation threshold, andwherein the processor is configured to locate the stationary radiationsource based on the reception time difference of the pairs of photons,and to generate the fault signal when the location of the stationaryradiation source exceeds the location deviation threshold.
 26. Theassembly of claim 25, wherein a pair of photons emitted by a positronannihilation event generates a positron emission path, wherein thethreshold parameter is a time difference range, and wherein theprocessor is configured to generate the fault signal when a differencebetween a reception time of the pairs of photons is outside of the timedifference range.
 27. An imaging assembly comprising: a gantrycomprising a first array of positron emission detectors and a secondarray of positron emission detectors opposing the first array ofdetectors; a housing coupled to the gantry, the housing comprising abore and an annular radiation source about the bore; and a processorconfigured to receive positron emission data from the first and secondarrays of positron emission detectors and to distinguish the positronemission data from the annular radiation source, and to generate a faultsignal when the positron emission data from the annular radiation sourceexceeds a threshold parameter.
 28. The assembly of claim 27, wherein theprocessor is further configured to concurrently classify the positronemission data from the annular radiation source and from a patient scanregion within the bore.
 29. The assembly of claim 28, wherein theprocessor is further configured with a spatial filter to distinguish thepositron emission data from the annular radiation source and from thepatient scan region.
 30. The assembly of claim 27, wherein the firstarray and second array of detectors are stationary.
 31. The assembly ofclaim 27, wherein the first array and second array of detectors arerotatable.
 32. An imaging assembly comprising: a gantry comprising afirst array of rotatable positron emission detectors and a second arrayof rotatable positron emission detectors opposing the first array ofdetectors; one or more radiation source holders coupled to the gantrysuch that the one or more radiation source holders are fixed relative tothe first array and the second array of detectors; and a processorconfigured to receive positron emission data from the first and secondarrays of rotatable positron emission detectors and to distinguish thepositron emission data from the one or more radiation source holders,and to generate a fault signal when the positron emission data from theone or more radiation source holders exceeds a threshold parameter. 33.The assembly of claim 32, wherein the gantry comprises a bore, whereinthe bore comprises a patient scan region spaced away from the one ormore radiation source holders, and the processor is further configuredto distinguish the positron emission data from the patient scan regionin the bore.
 34. The assembly of claim 31, wherein the one or moreradiation source holders comprise at least four radiation sourceholders.
 35. The assembly of claim 31, further comprising one or moreradiation sources held by the corresponding one or more radiation sourceholders, the one or more radiation sources comprising a radioactivity ofabout 1 μCi to 300 μCi and an energy of about 511 keV.
 36. The assemblyof claim 32, further comprising one or more radiation sources held bythe corresponding one or more radiation source holders, the one or moreradiation sources comprising a shape selected from the group consistingof a cylinder, sphere, and ring.
 37. An imaging method comprising:receiving concurrent positron emission data from a patient and acalibration source spaced away from the patient, using a first array ofpositron emission detectors and a second array of positron emissiondetectors opposing the first array of detectors; distinguishing thepositron emission data from the patient and from the calibration source;generating calibration data using the positron emission data from thecalibration source; generating patient data using the positron emissiondata from the patient; and generating a fault signal when thecalibration data exceeds a threshold parameter.
 38. The method of claim37, wherein distinguishing the positron emission data from the patientand from the calibration source comprises spatially filtering thepositron emission data.
 39. The method of claim 37, further comprisingadjusting the spatial filtering before applying the spatial filtering.40. The method of claim 39, wherein adjusting the spatial filtering isperformed using a patient treatment plan.
 41. The method of claim 38,wherein the spatial filtering of the positron emission data comprisesexcluding the positron emission data located outside a calibrationregion and a patient region.
 42. The method of claim 37, whereinreceiving the concurrent positron emission data occurs concurrently withgenerating the fault signal.
 43. The method of claim 37, furthercomprising treating the patient using a radiation source concurrentlywhile receiving the concurrent positron emission data from the patientand from the calibration source.
 44. The method of claim 43, furthercomprising stopping treatment of the patient using the radiation sourcein response to generating the fault signal.
 45. The method of claim 37,further comprising deactivating one or more of the positron emissiondetectors based on the generation of the fault signal.
 46. The method ofclaim 37, further comprising deactivating up to three of the first arrayand second array of detectors based on the generation of the faultsignal, wherein the fault signal comprises a fault in up to three of thedetectors.
 47. The method of claim 37, further comprising deactivatingall of the detectors based on the generation of the fault signal,wherein the fault signal comprises a fault in four or more of thedetectors.
 48. The method of claim 37, further comprising calibratingone or more positron emission detectors using the calibration data. 49.The method of claim 37, wherein the positron emission data correspondsto lines of response non-intersecting with a patient imaging field ofview of the detectors, the patient imaging field of view comprising apatient scan region.
 50. The method of claim 37, further comprisingverifying a positron emission detector calibration monitoring systemcoupled to the detectors based on the generation of the fault signal.